Recent advances in “bioartificial polymeric materials” based nanovectors

1.1 Bioartificial polymeric materials

“Bioartificial polymeric materials” (BPMs) are a class of polymeric composites based on the blend between synthetic and natural polymers, designed to produce new materials combining the biocompatibility of the biological components with the physical and mechanical features of the synthetics [1, 2, 3].These materials are engineered for different applications such as biodegradable delivery systems, leak proof membranes, systems of proteins purification, dialysis membranes, wound dressing, artificial skin, cardiovascular devices, nerve guide channels, implantable devices, bone graft substitutes [4, 56] and have an enormous repercussion in the human life quality [7]. Table 1summarizes the main applications of these materials (Table 1).

1.1.2 Synthetic polymers for BPMs

Synthetic polymers are petroleum-based products produced by chemical reactions. These materials are important components of BPMs thanks to their inert nature, high resistance of chemical linkages to hydrolytic and oxidative degradation and ability to tailor mechanical properties. Synthetic polymers contribute to the efficient functioning of devices providing mechanical support to implants such as articulating surfaces and scaffolds (e.g. knee and hip implants), protective coatings to improve blood compatibility, electrically stimulating devices (e.g. pacemakers, heart valves), catheters and dialysis tubing, vascular grafts and implantable drug delivery systems (e.g. drug eluting coatings on vascular stents). The main classes of synthetic polymers used in BPMs include poly(olefins), poly(urethanes), poly(carbonates), poly(siloxanes), poly(amides), poly(ethers), poly(sulfones) and poly(esters) [12, 13, 14, 15]. Table 3summarizes their chemical structures and general properties (Table 3).

Table 3:

Synthetic polymers used for bioartificial polymeric materials.

Polymer	Structure	Key properties
Poly(ethylene)	(C2H4)n	Excellent chemical resistance Low- and high-density grades Thermoplastics features
Poly(propylene)	(C3H6)n	Flexible with good fatigue resistance Thermoplastics features
Poly(methyl methacrylate)	(C5O2H8)n	Good impact strength Lightweight, transparent Poor chemical resistance Thermoplastics features
Poly(dimethyl siloxane)		Excellent viscoelastic properties Transparent, elastic, inert, nontoxic
Poly(ether ether ketone)		Excellent chemical resistance Excellent mechanic properties Thermoplastics features
Polyurethane		Biodegradable Biostable Thermoplastics features

1.2 Nanotechnology and medicine

Nanotechnology (NT) is the science of manipulating matter at the atomic or molecular scale and holds the promise of providing significant improvements in the technologies intended to enhance human well-being and protect the environment [15]. NT is often regarded as a product of the latter part of the twentieth century but it influenced human evolution from the earliest civilization. Indeed, the ancient Greeks used permanent hair-dying recipes composed of 5 nm lead sulfide crystals and European medieval artists colored stained glass using metal nanoparticles. Modern NT, started in 1959 when physicist Richard Feynman recognized the possibility to build machines able to manufacture objects with atomic precision and explained that, at the nanoscale, surface phenomena dominate the behavior of the objects [16]. The term NT, however, was introduced in 1974 by Norio Taniguchi referring to the “production technology to get the extra-high accuracy and ultra-fine dimensions” [17]. Practical application of NT started with the description of the molecular manufacturing [18] and the invention of the scanning tunneling microscope (STM) that allowed the first direct manipulation of individual atoms [17]. Nowadays, NT is a dynamic field where over 50,000 articles published annually and more than 2,500 patents filed [19].

Nanomedicine, an offshoot of NT, uses nano-sized tools for the diagnosis, prevention and treatment of diseases (Figure 1). Applicative examples are biosensors, implantable devices, prostheses components and drug delivery platforms. This chapter focuses on the delivery of therapeutic substances through bioartificial polymeric nanovectors, a novel and interesting aspect of nanomedicine.

Figure 1:

Nanovectors direction to the FDA approval.

2.1 Nanospheres and Nanocapsules

Nanospheres (NSs) and nanocapsules (NCs) are nano-sized vectors composed of amphiphilic copolymers structured with hydrophobic chains forming the inner part of the particles and hydrophilic portions on the surface. NSs have homogeneous solid matrices [39] while NCs exhibit a core-shell structure in which the drug is confined to a reservoir or within a cavity surrounded by a polymer membrane [40]. Both NSs and NCs allow the fine tuning of their properties through surface functionalization, the use of different shell materials and with the regulation of their size [41, 42, 43]. Poly-lactic acid (PLA), poly-glycolic acid (PGA), poly-lactic-co-glycolic acid (PLGA), poly ε-caprolactone (PCL), chitosan (CS) and polyethylene glycol (PEG) are the main materials used for the synthesis of these systems due to their wide biocompatibility and biodegradability [44, 45, 46, 47, 48]. NS drugs are dissolved, entrapped, encapsulated, chemically bound or adsorbed to the constituent polymer matrix [49, 50] while NCs carry the active substance in the core, or on their surfaces or absorbed in the polymeric membrane [51, 52, 53]. The use of NSs and NCs is an attractive strategy for the vectorization of a variety of active substances such as antineoplastics, antiinflammatories, immunosuppressants, antigens, hormones, antivirals, antibacterials, antifungals, diuretics, antipneumocystics and vitamins. Moreover, these systems are useful to mask unpleasant tastes, to provide controlled release properties, to protect vulnerable molecules from degradation and to increase the therapeutic efficacy of active molecules [54, 55]. Finally, NSs and NCs have high intracellular uptake and require a low amount of polymer for each particle, resulting in high drug loading [46, 49]. NSs and NCs have common preparation techniques that are classified into two general categories depending on the starting material. The use of monomers requires emulsion polymerization, interfacial polymerization or ionic gelation methods. Differently, for preformed polymers, nanoparticle preparation is achieved through emulsification/solvent evaporation, emulsification/solvent diffusion, salting out, dialysis and nanoprecipitation. The emulsion polymerization method is carried out using organic or aqueous solvents as continuous phase [46]. Surfactants or protective soluble polymers are used to prevent aggregation in the early stages of the polymerization. The polymerization process starts using an initiator molecule (ion or free radical), or activating the monomer with high-energy radiations. Incorporation of active principles is obtained dissolving the substance in the same phase of the monomers. The interfacial polymerization is a process with a similar mechanism that involves the dissolution of two reactive agents into two phases (i.e., continuous- and dispersed-phase), with the reaction that takes place at the interface of the two liquids [56]. Interfacial polymerization permits to modulate the formation of NSs or NCs using different eluents. In fact, to promote NC formation, aprotic solvents are used, while protic liquids induce the formation of NSs [57]. Incorporation of active principles is obtained dissolving the drug into the dispersed phase. Ionic gelation permits the preparation of polymeric nanoparticles using biodegradable hydrophilic polymers such as CS, gelatin and sodium alginate. This method requires the mixture of two aqueous phases, one containing the hydrophilic polymer and the other a crosslinker (e.g. poly-anion sodium tripolyphosphate). The positive groups of the polymer interact with negative charged crosslinkers to form nano-sized vectors [58]. Drug is added in the same phase of the hydrophilic polymer [58]. Emulsification/solvent evaporation is a method that requires the preparation of the polymer solution in lipophilic volatile solvent with a subsequent formation of an emulsion, adding water and stabilizers. The lipophilic solvent diffuses through the emulsion and its evaporation lead to the formation of a nanoparticle suspension. High-speed homogenization or sonication is utilized to improve the diffusion, while the solvent evaporation is favored by continuous magnetic stirring at room temperature or under reduced pressure. The solidified nanoparticles are collected by centrifugation and washed with distilled water to remove additives [59]. Drug loading is carried dispersing the substance in the volatile solvent. Similarly, in the emulsification/solvent diffusion method, the encapsulating polymer is dissolved in a partially water soluble eluent and saturated with water to ensure the diffusion. Subsequently, the polymer-water saturated solvent phase is emulsified in an aqueous solution containing stabilizers, leading to the solvent diffusion to the external phase and to the formation of nanovectors. Drug loading is achieved dissolving the substance in the polymer phase [46]. Salting out is a modification of the emulsification/solvent diffusion in which polymer and drug are initially solubilized in the volatile solvent which is emulsified into an aqueous gel containing salting-out agents (e.g. electrolytes, such as magnesium chloride, calcium chloride, and magnesium acetate, or non-electrolytes such as sucrose) and colloidal stabilizers (e.g. polyvinylpyrrolidone or hydroxyethylcellulose). This oil/water emulsion is diluted with aqueous solutions to enhance the diffusion. The salting out agents improve the encapsulation efficiency of the drug [46]. In the dialysis methodology, the polymer is dissolved in an organic solvent and placed inside a dialysis tube. Dialysis is performed against a non-solvent miscible with the lipophilic eluent. The progressive aggregation of polymer and the formation of nanoparticles is the consequence of the displacement of the organic solvent inside the membrane [59]. Drug incorporation is obtained adding the active principle in the same eluent of the polymer. Finally, in the nanoprecipitation method, the polymer is solubilized in a water-miscible solvent and is injected into a stirred aqueous solution containing a surfactant. The stirring causes a fast diffusion of the solvent and the polymer deposition on the interface between the water and the organic eluent, leading to the instantaneous formation of a colloidal suspension [60]. The aqueous solution must be a non-solvent of the polymer [61]. Drug encapsulation is achieved solubilizing the drug into the organic solvent [61]. A schematic representation of the described preparation techniques is available in Figure 2. Bioartificial polymers are widely studied as components of NSs and NCs. Such systems are described by Bellotti et al. as composed of butyl methacrylate, poly(ethylene glycol) methyl ether methacrylate, 2-(dimethylamino) ethyl methacrylate crosslinked with trimethylolpropane trimethacrylate and functionalized with folic acid on their surface in order to specific target enclosed anticancer drug to cancer cells [62]. The antitumor activity is the subject of research of several other authors. For example, Cui et al. formed ionically assembled nanoparticles from poly(ionic liquid-co-N-isopropylacrylamide) with deoxycholic acid through electrostatic interactions. These nanoparticles exhibit dual-responsive properties based on pH and thermal environment conditions with practical applications as drug delivery carriers, as shown by the encapsulation of doxorubicin. In particular, low pH and high temperature provoke structural collapse of the ionically assembled nanoparticle and the release of doxorubicin. In fact, 80 % of drug molecules are released within 48 h at pH 5.2, 43 °C, but only 30 % of doxorubicin is released within 48 h at 37 °C and pH 7.4 [63]. Bahadur et al. designed nanoparticles formed by poly(2-(pyridin-2-yldisulfanyl)ethyl acrylate) conjugated with PEG and cyclo(Arg-Gly-Asp-d-Phe-Cys) peptide. These nanovectors are loaded with doxorubicin. The size of the vehicle is 50.13 ± 0.5 nm in PBS. Such vectors are stable in physiological condition and release doxorubicin with the trigger of acidic pH and redox potential. Moreover, these acrylate-based nanoparticles show a two-phase release kinetics, providing both loading and maintenance doses for cancer therapy. The conjugation with the peptide enhances the cellular uptake and nuclear localization. In fact, these vectors exhibit significantly higher anticancer efficacy compared to that of free doxorubicin at concentrations higher than 5 μM [64]. Barick et al. synthesized glycine functionalized magnetite (Fe₃O₄) nanoparticles by Michael addition/amidation reaction. These nanocarriers have average size of about 10 nm and are resistant to protein adsorption in physiological medium. Moreover, the terminal amino acids on the shell of the magnetic nanocarriers allow outer functionalization and potential conjugation with drug molecules. The encapsulation of doxorubicin as model drug revealed high loading affinity, sustained release profile, magnetic-field-induced heating and substantial cellular internalization. Moreover, the enhanced toxicity to tumor cells using a local magnetic field suggests their potential for combination therapy involving hyperthermia and chemotherapy [65]. Similarly, Zhao et al. produced arginine–glycine–aspartic acid-modified Fe₃O₄ nanoparticles to control the delivery and release of doxorubicin. The conjugation of these targeted magnetite nanoparticles with the drug is via acid-labile imine bond. Such linkage gives magnetic control, specific targeting and pH-responsivity to the nanocarriers. The cell toxicity assays indicate higher anticancer activity of these pH-sensitive magnetic nanocarriers compared to free doxorubicin and increased cytotoxicity consequent to the conjugation with arginine–glycine–aspartic acid peptides [66]. Cheng et al. developed nanoparticles of carboxy-terminated poly(d,L-lactide-co-glycolide)-block-poly(ethylene glycol) conjugated with A10 RNA aptamers, able to bind the prostate specific membrane antigens. Such nanoparticles deliver docetaxel and paclitaxel to tumor cells. These nanovectors are evaluated in a xenograft mouse model of prostate cancer. The surface functionalization with A10 aptamers significantly enhances the delivery to tumors [67]. Patil et al. synthesized copolymer PLA–PEG nanoparticles functionalized with biotin or folic acid and incorporating paclitaxel, by solvent polymerization technique. The addiction of the ligands significantly enhances nanoparticles accumulation in tumor cells in vitro and results in improved efficacy of in a mouse xenograft tumor model [68]. Farokhzad et al. synthesized a bioconjugate composed of PLA-block-PEG copolymer and aptamers for targeted delivery to prostate cancer cells. These nanovectors encapsulate the model drug rhodamine labeled with dextran. Such nanoparticles present carboxylic acid groups on the particle surface, useful for functionalization and for covalent conjugation with amine-modified aptamers. Moreover, the coating of PEG enhances circulating half-life and decreases the uptake into non-targeted cells. The bioconjugation with RNA aptamers permits the targeting on prostate LNCaP epithelial cells [69]. Schiffelers et al. produced self-assembling nanoparticles with siRNA and polyethyleneimine PEGylated with an Arg-Gly-Asp (RGD) peptide ligand attached at the distal end of the PEG. These nanovectors deliver siRNA inhibiting vascular endothelial growth factor receptor-2 expression into tumor neovasculature expressing integrins. Intravenous administration of this system into tumor-bearing mice results in selective tumor uptake, siRNA sequence-specific inhibition of protein expression within the tumor and reduction of both tumor angiogenesis and growth rate [70]. Cho et al. synthesized retinoic acid loaded poly(L-lactic acid) nanoparticles coated with galactose-carrying polymer for hepatocyte-specific targeting using galactose ligands as recognition signals to asialoglycoprotein receptors. The authors study the effects of released retinoic acid on morphology and DNA synthesis of hepatocytes. Such drugs modify in vitro shapes of hepatocytes. Moreover, fluorescence and confocal laser microscopic studies confirm the positive influence of galactose-carrying polymers coating on nanoparticles internalization [71] Soppimath et al. synthesized core-shell nanoparticles, self-assembled from the amphiphilic tercopolymer poly(N-isopropylacrylamide-co-N,N-dimethylacrylamide-co-10-undecenoic acid) in which 10-undecenoic acid is employed as hydrophobic and pH-sensitive segment. The temperature responsiveness of the core-shell nanoparticles is triggered by a change in the environmental pH. The shell of these nanoparticles is composed of amine groups able to conjugate biological signals for specific affinities to certain cell types. Such nanoparticles, loaded with doxorubicin, are stable in PBS at 37 °C but precipitate in acidic environment, triggering the release of the enclosed drug molecules [72]. Shu et al. produced crosslinked hollow polyelectrolyte NCs composed of cysteamine conjugated CS and dextran sulfate by adsorption on β-cyclodextrin functionalized silica spheres. These NCs have enhanced physical stability against acidic pH conditions and decrease the loss of protein caused by the gastric cavity and the release of drugs in the intracellular environment after glutathione reduction. Bovine serum albumin (BSA) used as model drug exhibits spherical morphology, dimension of 120 nm, with a good polydispersion index and sustained release without the initial burst [73]. The cited vehicles are summarized in Table 5.

Figure 2:

Schematic representation and manufacturing methods of nanospheres –left- and nanocapsules –right.

2.2 Nanohydrogels and nanoaggregates

Nanohydrogels (NHYs) are nano-sized networks of polymer chains able to incorporate H₂O in their structure. Usually, NHYs are macromolecular hydrocolloids with numerous hydrophilic functional groups [74]. Their composition ranges from linear water-soluble polymers to water insoluble molecules that act as swellable networks stabilized by crosslinking agents. In general, these substances have high molecular weight (5000–10000 Da) and cannot cross-biological membranes. Further, they include cellulosic components like sodium carboxymethyl cellulose or polyanion bioadhesives like polyacrylic acid. These nanovectors are classified on the basis of the presence or absence of electrical charge located on the crosslinked chains. In fact, they are grouped as nonionic, anionic, cationic, amphoteric electrolytes or zwitterions. The surface charge regulates the adhesivity. For example NHYs, due to their capability of forming strong non-covalent bonds with the mucin, have prolonged ocular residence time and reduced dosing frequency [75]. Production of hydrogels requires the simple preparation of polymer solutions in low or intermediate concentrations and the formation of crosslinks for the prevention of the dissolution. Many crosslinking methods are currently available for hydrogel synthesis. Generally, physically crosslinked gels are those whereas physical interactions exist between polymer chains (e.g. hydrogen bonds, amphiphilic graft) while chemically crosslinked hydrogels are synthesized with covalent bonds (e.g. crosslink with aldehydes, free radical polymerization). The nanodimensions are usually obtained with sonication [76]. The medical application of nano-sized hydrogels is limited by the difficult administration of an accurate dose of active principle due to the variable release of gelified systems [75]. Nanoaggregates (NAGs) are colloidal carriers formed from amphiphilic block copolymers. In some cases, further molecules act as crosslinker agents. NAGs possess inherent properties such as high loading efficiency and in vivo stability. These vehicles are able to provide site-specific drug delivery via either a passive or active targeting mechanisms. NAGs are suitable for encapsulation of poorly water-soluble drugs by covalent conjugation as well as physical encapsulation. Active transport is achieved by conjugating a drug with vectors or ligands that bind specific receptors [77]. The synthesis of NHYs and NAGs is summarized in Figure 3. The use of bioartificial materials for the preparation of NHYs and NAGs is a novel research interest. Despite of this, a number of papers are available in literature. For example, CS-poly (acrylamide-co-methacrylic acid) hydrogels were synthesized by Ullah et al. They use different coupling agents (3-dimethylaminopropyl)-3-ethylcarbodimide hydrochloride and 3-aminopropyltriethoxysilane) and the functionalization with phenylboronic acid, a glucose sensing moiety, to design multifunctional NHYs with enhanced glucose sensitivity, stability, drug loading and release profile. Moreover, the authors study the glucose-induced volume phase transition and release profile at physiological conditions of the model drug Alizarin Red (a compound with 1,2-diol structure, similar to insulin) in order to find potential application in self-regulated insulin delivery with enhanced sensitivity toward glucose [78]. Jaiswal et al. synthesized poly N-isopropylacrylamide – CS-based NHYs encapsulating iron oxide (Fe₃O₄) magnetic nanoparticles through free radical polymerization of the acrylate in presence of CS. These NHSs are spherical shaped with size ranging from 50 nm to 200 nm on the base of the feed ratios of CS. The encapsulation of Fe₃O₄ nanoparticles into poly N-isopropylacrylamide–CS based NHYs is confirmed by transmission electron microscopy. This system shows optimal magnetization, good specific absorption rate and excellent cytocompatibility, finding potential applications in hyperthermia treatment of cancer and targeted drug delivery [79]. Yoon et al. produced self-assembled NAGs co-encapsulating doxorubicin and oligonucleotides through the conjugation of four-arm poly(ethylene glycol) with doxorubicin and anti-bcl-2 oligonucleotides. These conjugates are hydrophobically self-assembled into NAGs in aqueous solutions. Elemental scanning of the products reveals a core-shell structure with the drug located at the core of the vectors and the genetic materials at the shell. Analysis by dynamic light scattering and electron microscopy proves the complete disappearance of the particles under reducing conditions and the liberation of oligonucleotides at low pH. In vitro studies confirm the uptake of drug and oligonucleotides in cells treated with NAGs [80]. Table 6recaps the described nanovectors (Table 6).

Figure 3:

Schematic representation and manufacturing methods of nanoaggregates –top- and nanohydrogels –bottom.

2.3 Micelles (MCs) and solid lipid nanoparticles (SLNs)

Micelles (MCs) are colloidal dispersions belonging to a large family of systems consisting of particulate matter (the “dispersed phase”), distributed within a continuous phase (the “dispersion medium”), usually constituted by water. MCs form spontaneously at certain concentration (critical micelle concentration) and temperature (critical micelle temperature) values from amphiphilic or surface active agents. Usually, these vehicles have particle size ranging between 5 and 100 nm [81, 82]. Regarding the structure, the hydrophobic fragments of amphiphilic molecules form the core of MCs, while hydrophilic fragments the shells. The formation of MCs is driven by the decrease of free energy in the system because of the removal of hydrophobic fragments from the continuous phase, and the re-establishing of hydrogen bond network in water. Moreover, additional energy results from the formation of van der Waals bonds between hydrophobic blocks in the core of the formed MCs [83]. MCs possess high stability both in vitro and in vivo, good biocompatibility and are able to solubilize a broad variety of poorly soluble pharmaceuticals through the interaction of the lipophilic substances with their hydrophobic core [84]. Many of these drug-loaded vehicles are currently at different stages of preclinical and clinical trials [85]. Micellar nano-drug delivery systems have increased water solubility, improved bioavailability, reduction of toxicity, enhanced permeability across the physiological barriers, substantial changes in drug biodistribution, extended blood half-life and protection from degradation [86]. Moreover, MCs have spontaneous interstitial penetration into the body compartments with leaky vasculature (tumors and infarcts) [87]. Active targeting of MCs is obtained through surface chemical attachment of driving molecules [88]. MCs are prepared simply dissolving the amphiphiles in water. These vectors are thermodynamically stabilized against disassembly if the amphiphilic concentration remains above the CMC. While, upon dilution below the CMC, MCs disassemble with a rate depending on the structure of the amphiphiles and on the interactions between the chains [89]. The encapsulation of molecules is obtained dissolving the substance into the micellar solution [89]. A further method of MCs preparation, useful to encapsulate non-water soluble molecules, is the thin-film hydration method. The amphiphilic copolymer and the lipophilic drug are dissolved in organic solvent. Then, the apolar eluent is removed to get the drug-containing lipid membrane. This film is resuspended in a polar solvent for nano-MCs self-assembly. The driving force of this process is the hydrophobic effect between the non-polar segments of the polymers. The hydrophobic effect also plays an important role in the drug encapsulation, stabilizing the intermolecular interaction between the substance and the hydrophobic segment [90]. SLNs are colloidal carrier systems, generally spherical in shape, composed of a high melting point lipids, as a solid core, coated by aqueous surfactants. The core lipids are fatty acids, acylglycerols and waxes, whereas phospholipids, sphingomyelins, bile salts and sterols are utilized as stabilizers [91]. SLNs are useful for the delivery of poor water soluble drugs [92]. The particle diameters are in the range of 10–1000 nm. SLNs are characterized by high biocompatibility, high bioavailability, physical stability, protection of incorporated labile drugs from degradation, excellent tolerability, prevention of problems related with multiple routs of administration, avoidance of the use of organic solvents during the preparation, formation of films over the skin showing occlusive properties and absence of problems concerning large-scale production and sterilization [93, 94]. However, common disadvantages of SLNs are particle growth, unpredictable gelation tendency, uncertain diffusion of the drug within the lipid matrix of the vector, unexpected dynamics of polymorphic transitions and inherent low incorporation rate due to the crystalline structure of the solid lipid [95, 96]. SLNs are prepared through homogenization, solvent-evaporation, microemulsion and film-ultrasound dispersion techniques. In the homogenization method, the homogenizers push a liquid with high pressure (100–2000 bar) through a narrow gap. The high shear stress and cavitation forces disrupt the particles down to the submicron range. This technique is carried in hot or cold conditions. Hot homogenization requires temperatures above the melting point of the lipid and necessitates the preparation of a pre-emulsion of the drug loaded lipid melt with the aqueous emulsifier. Higher temperatures result in lower particle sizes. In cold conditions, the drug containing lipid melt is cooled and dispersed into a cold surfactant solution. This pre-suspension is homogenized at or below room temperature, breaking the lipids in solid nanoparticles. In the solvent-evaporation method, the lipids are dissolved in a water-immiscible organic solvent that is emulsified in an aqueous solvent. The nanoparticles dispersion is obtained upon evaporation of the eluent that leads to lipid precipitation. The microemulsion technique is operated stirring a mixture of low melting fatty acids, emulsifiers and water, at 65–70 °C. This hot liquid is dispersed in cold water (2–3 °C) under stirring. The high-temperature gradient facilitates rapid lipid crystallization and prevent aggregation. Finally, in the film-ultrasound dispersion method, the lipid and the drug are put into suitable organic solution that is evaporated to form a lipid film. The following addition of an aqueous solvent results in an emulsion that is sonicated giving SLNs with uniform particle size [97]. A recap of the production methods for both MCs and SLNs is available in Figure 4. Bioartificial polymers are widely used in the preparation of both these vesicular systems. For example, Zhang et al. used the core crosslinking method to generate MCs able to increase curcumin delivery to HeLa cells (immortalized cancer cells) in vitro and improve tumor accumulation in vivo. These MCs are designed with folic acid-PEG as the hydrophilic unit, pyridyldisulfide as the crosslinkable and hydrophobic unit, and disulfide bond as the crosslinker. Such nanovectors show spherical shape with a diameter of 91.2 nm and high encapsulation efficiency. Cytotoxicity effectiveness is demonstrated by the high cellular uptake and the positive in vitro antitumor studies. The linkage with folate targets the curcumin against cancer cells and enhances the in vivo efficacy of these MCs [98]. Similarly, Lee et al. produced pH-sensitive polymeric MCs composed of poly(l-histidine), PEG and poly(L-lactic acid) block copolymers with folate conjugation, delivering adriamycin. These MCs are investigated for pH-dependent drug release, folate receptor-mediated internalization and cytotoxicity using MCF-7 cells (human breast adenocarcinoma‎ cell line) in vitro. These nanovectors show accelerated drug release only at acidic pH. Moreover, the conjugation with folic acid enhances tumor cell kill due to folate receptor-mediated tumor uptake [99]. Li et al. reported linear PEG and dendritic cholic acids block copolymers MCs stabilized with boronate esters at the core–shell interface for efficient anticancer drug delivery. Such system is loaded with paclitaxel to assess its capacity to retain the encapsulated drug under physiological conditions and release the payload when triggered by the lower pH value of the tumor environment or by the presence of competitive diols (e.g. mannitol). This nanovector shows minimal premature drug release at physiological glucose level and physiological pH values in blood circulation and simple activation at the acidic tumor microenvironment or by the additional intravenous administration of mannitol as an on-demand triggering agent [100]. Paclitaxel-loaded mixed polymeric MCs consisting of poly(ethylene glycol) distearoyl phosphoethanolamine conjugates, solid triglycerides and cationic lipofectin lipids were prepared by Wang et al. Optimized MCs have average size of about 100 nm, and zeta-potential of about −6 mV. Such vehicles are stable when stored at 4°C or at room temperature. Release of paclitaxel starts at 37 °C and, approximately, 16 % of the drug is dispensed in 72 h. In vitro anticancer effects of the nanovectors are evaluated using human mammary adenocarcinoma (BT-20) and human ovarian carcinoma (A2780) cell lines. The results show enhanced anti-cancer activity due to the ability of the MCs to escape from endosomes and enter the cytoplasm of BT-20 and A2780 cancer cells [101]. Lee et al. developed paclitaxel-loaded sterically stabilized SLNs for parenteral administration. These nanovectors, prepared using the hot homogenization method, are composed of trimyristin as a solid lipid core and egg phosphatidylcholine and pegylated phospholipid as stabilizers. The particles are spherical in shape, with sizes and zeta potentials of around 200 nm and −38 mV. Paclitaxel is loaded to the solid cores at a w/w ratio of 6 % with high encapsulation efficiency. In vitro drug release studies show a slow sustained release and high cytotoxicity on OVCAR-3 human ovarian cancer cell line and MCF-7 breast cancer cell line [102]. Gao et al. prepared poly(ethylene glycol)/phosphatidyl ethanolamine (PEG−PE) conjugates for the solubilization and delivery of various poorly soluble anticancer drugs such as m-porphyrin, tamoxifen and taxol. These MCs are stable and have the size of 10 to 40 nm [103]. Wong et al. investigated the in vivo efficacy, unwanted toxicity and loco-regional distribution of doxorubicin-loaded polymer-lipid hybrid nanoparticles formulation in a murine solid tumor model after intratumoral injection. These SLNs are prepared by dispersing the drug in stearic acid and tristearin, with subsequent addition of the hydrolyzed polymers of epoxidized soybean oil to enhance doxorubicin incorporation into the lipids. This formulation is injected intratumorally in murine solid tumors of approximately 0.3 g. In vivo, SLNs-treated tumors develop substantially larger central necrotic regions compared to the untreated tumors, with minimal systemic toxicity [104]. Kukowska-Latallo et al. produced polyamidoamine dendritic polymers conjugated to folic acid as targeting agent for methotrexate. These conjugates are injected intravenous into immunodeficient mice bearing human KB tumors overexpressing the folic acid receptors, resulting in high internalization into the tumor cells [105]. Table 7 acts as a recap of the described nanosystems.

Figure 4:

Schematic representation and manufacturing methods of micelles –left- and solid lipid nanoparticles –right.

2.4 Nanofibers

Nanofibers (NFs) are fibers with diameters less than 100 nanometers that exhibit special properties due to the extremely high surface to weight ratio, low density, high pore volume and tight pore size [106]. These properties make NFs suitable for applications ranging from medical (e.g. drug delivery systems) to industrial and high-tech fields (e.g. aerospace, capacitors, transistors, battery separators, energy storage, fuel cells) [106]. In nanomedicine, NFs are widely used due to their similarity to the extracellular matrix (ECM), the possibility to use several materials for their synthesis (in fact, natural and synthetic polymers along with several solvent systems are effectively used to create NFs) and the possibility to change their architecture in regard to porosity, diameter, mechanical properties, structure arrangement and structure functionalization [107]. The synthetic techniques for NFs are self-assembly, phase separation and electrospinning. All of these techniques require the preparation of a homogeneous drug–polymer solution that is loaded in the electrospinning machine (electrospinning technique), is simply mixed (self-assembly method) or is treated to obtain gelation and freeze-drying (phase separation) (Figure 5) [106, 107]. Research about NFs as drug delivery systems is at the early stage of exploration and most of the works focus on the sustained release profiles of model drugs (e.g. small molecules, herbs, proteins, DNA, genes and vaccines) using biodegradable hydrophilic, hydrophobic or amphiphilic polymers and, recently, BPMs [108]. Zhang et al. reported degradable heparin-poly (ε-caprolactone) fiber mats fabricated by electrospinning. The highly sulfated heparin heteropolymer remains homogenous in the spinning solution and is distributed throughout the fabricated polymers. The NFs release heparin for 14 days with a diffusionally controlled kinetics over this period. The drug retains its biological properties and functionality [109]. Chew et al. investigated the encapsulation of human β-nerve growth factor (NGF), stabilized in BSA carrier proteins, in a copolymer of ε-caprolactone and ethyl ethylene phosphate. The proteins are randomly dispersed throughout the electrospun fibrous mesh in an aggregated form. The sustained release of NGF by diffusion is detectable for at least 3 months and the bioactivity of the drug is retained throughout this period [110]. Feng et al. produced CS polyethylene oxide NFs with uniform diameter of 112 nm and the potential modulation of CS viscosity and surface tension through the use of different CS molecular weight and polyethylene oxide quantities. These vehicles exhibit excellent biocompatibility with hepatocytes [111]. Similarly, Bhattarai et al. reported that CS\polyethylene oxide nanofibrous scaffolds promote the attachment of human osteoblast and chondrocytes, maintaining their characteristic morphology and viability. This matrix is of particular interest for tissue engineering, drug delivery and tissue remodeling [112]. Moreover, Subramanyan et al. prepared CS\polyethylene oxide NFs for cartilage tissue engineering. These scaffolds are used for cell attach and deliver of growth factors [113]. Park et al. produced chitin/poly glycolic acid NFs with BSA coating to improve human epidermal fibroblasts attach and spread [114]. Shalumon et al. developed bioactive and biocompatible NFs composed of carboxymethyl cellulose\polyvinyl alcohol blend. Such nanomaterials are tested for cytotoxicity and cell attachment, resulting in a safe application for tissue engineering and drug delivery [115]. Nanofibrous scaffold of CS\polyvinyl alcohol and carboxyethylchitosan\polyvinyl alcohol are also prepared by Zhou et al. These materials, tested on L929 fibroblast culture, have good cell attachment and growth [116]. CS hydroxyapatite nanofibrous scaffolds are reported by Yang et al. This nanomaterial significantly stimulates the bone forming ability due to the excellent osteoconductivity of hydroxyapatite [117]. Finally, Junkasem et al. described the fabrication of α-chitin whisker-reinforced poly(vinyl alcohol) NFs by electrospinning. The α-chitin whiskers are prepared from α-chitin flakes by acid hydrolysis. Such vectors exhibit average length and width of about 549 and 31 nm, respectively. The incorporation of the chitin whiskers within the poly(vinyl alcohol) is verified by infrared spectroscopy and thermogravimetry, resulting in an increased Young’s modulus of the bioartificial polymer of about 4–8 times compared to the unmodified poly(vinyl alcohol) [118]. Table 8summarizes the described bioartificial polymeric NFs (Table 8).

Figure 5:

Schematic representation and manufacturing methods of nanofibers.