Polymeric NPs represent a highly effective nanoplatform for drug delivery [8]. An attractive feature of polymer-based drug delivery platforms is their controlled-release properties which is attainable via the use of biodegradable monomers [9]. Polymer NPs are capable of drug encapsulation and release in a temporally controlled manner, a property which can be achieved through surface or bulk erosion of the polymeric NPs, diffusion of drugs out of the polymer mesh network, or swelling of the polymer matrix and subsequent diffusion [10]. Drug release in polymeric systems may also be achieved in a triggered or smart manner in addition to temporal release, which can be manifested by response of the polymers to various triggers in their environment such as pH, heat, enzymes, or exposure to other appropriate sources of energy or chemicals that can break bonds or displace the drug molecules [10]. Their property of longitudinal controlled-drug release renders polymeric nanomedicines highly appealing as drug delivery vehicles, since drug release can be optimally achieved at the site of action for over long periods, without the patient requiring repeat dosing. Examples of drug delivery and imaging applications of polymeric NPs will be discussed in the subsequent sections of this chapter.

Oncology is one of the areas where nanomedicine has had the most impact to date, in particular with many NPs developed for the treatment and imaging of PCa [11, 12]. Both actively and passively targeted nanomedicines have been extensively utilized for oncology applications, in particular for drug delivery to solid tumors. If drug delivery NPs are to successfully reach the tumor and effectively deliver their drug load to cancer cells, they must be capable of effectively maneuvering the complex in vivo environment, and in particular avoiding clearance by the host phagocytic cells post-systemic administration. Indeed the dose of NPs that reach the tumor is dependent on a number of biophysicochemical properties of the NPs. These include the chemical signatures of the constituent material, be it lipids, polymers, or other inorganic components; the nature of the encapsulated drugs; the size and shape of the NP; and surface charge and hydrophilicity. In addition to these NP parameters, the tumor microenvironment, such as the degree of vascularization, necrosis, and size can also be factors influencing the navigation of NPs within tumors and hence their therapeutic efficacy. In the case of nontargeted NP-mediated drug delivery, the non-heterogeneous nature of the tumor microenvironment and the presence of abnormal leaky vasculature, termed the enhanced permeability and retention (EPR) effect, are beneficial as NPs can effectively extravasate through these distorted blood vessels and accumulate in tumor tissues at high concentrations. The degree of extravasation is inversely proportional to NP size, and smaller particles with sizes  <  150 nm are deemed to be most effective at trans-endothelial passage into tumor tissue [13].

Considering that docetaxel (Dtxl) is the common first-line chemotherapeutic treatment in castration-resistant PCa and given the adverse effects of Taxotere, which is the clinically approved formulation of Dtxl, Cervin et al. formulated liquid crystal nanoparticles (LCNPs) and showed these NPs to be effective at inhibiting PC3-induced tumors in SCID mice [14]. These 80–90 nm NPs were formulated with the lipids phosphatidyl choline, glycerol dioleate, and polysorbate 80, and their efficacy on tumor growth post-systemic administration was compared to that of Taxotere and empty LCNPs. In one study, the results from these nontargeted and passively accumulating NPs showed that the LCNP/Dtxl formulation led to the highest level of tumor regression of volumes decreasing to 10% or less in comparison to 18% and 70% for Taxotere, 10 days post-administration (total Dtxl dose of 1.62 mg used for entire course of study) [14].

The EPR effect can be successfully utilized for drug delivery to solid tumors, and other disease indications that lead to impaired lymphatic drainage and aberrant vasculature. However, active targeting using affinity ligands may lead to more efficient retention of nanomedicines within tumor sites [15].

Targeted NPs have the advantage of faster retention within tumors due to selective binding to overexpressed receptors on the surface of cancer cells. It is envisaged that the development of targeted NPs which represent the next generation of nanomedicines entering the clinic will lead to more effective therapeutic and imaging agents for PCa [16–18]. These vehicles can be engineered to recognize biophysical characteristics that are unique to the cancer cells. Most commonly, this recognition process involves the binding of vehicles to antigens that are expressed on the plasma membrane of the targeted cells. Small molecule ligands, antibodies, antibody fragments, and peptides have been used as targeting moieties in targeted nanoparticle development. A particular class of targeting ligands with high affinity to PCa cells is aptamers (Apts). Aptamers are single-stranded DNA or RNA oligonucleotides that fold into well-defined 3D structures that are capable of high affinity toward proteins, phospholipids, sugars, and nucleic acids [19]. Aptamers are ideal targeting ligands as they are non-immunogenic and exhibit remarkable stability in a wide range of pH (4–9), temperature, and organic solvents, with minimal loss of activity [20]. Furthermore, Apts can be chemically synthesized without the need for biological production, leading to reduction in batch-to-batch variability, which is one major advantage of Apts in comparison to larger and bulkier antibodies that can elicit toxic immune responses and are not easily produced on large scales [13].

We have shown that the generation of NP–Apt bioconjugates with nuclease-stabilized A10 2-fluoropyrimidine RNA Apts that bind to the prostate-specific membrane antigen (PSMA) efficiently targets and enters prostate epithelial cells which express the PSMA protein [21]. PSMA is a well-known characterized transmembrane protein that is overexpressed on PCa epithelial cells, is involved in membrane recycling, and becomes internalized through ligand-induced endocytosis [22]. Using the FDA approved biodegradable drug delivery polymer poly(d,l-lactic-co-glycolic acid) (PLGA) as a controlled release polymer, we were able to develop NPs with surfaces decorated with Apts that were targeted to the PSMA on the surface of PCa cells [20]. The developed Apt-targeted polymer NPs demonstrated a 77-fold increase in binding to prostate LNCaP cells in comparison to nontargeted polymeric NPs. To protect these particles from macrophages and increase their circulation time in vivo, the particles were also functionalized with the widely used poly(ethylene glycol) (PEG) polymer. These proof-of-principle studies using Apts as targeting ligands led us to develop and pioneer Apt-conjugated polymeric NPs to improve the therapeutic index of drugs by targeted delivery and controlled release to PCa cells [21, 23]. To achieve differential cytotoxicity against PCa cells, Apt-conjugated polymeric NPs were loaded with Dtxl, which has been demonstrated to prolong survival of patients with hormone-resistant PCa [20]. Figure 15.1 presents the general procedure used to prepare Dtxl-encapsulated pEGylated PLGA NP–Apt bio-conjugates, which were formulated using a nanoprecipitation technique. The targeted polymeric NPs were formulated by first co-preci-pitating Dtxl with the diblock polymer poly(lactide-co-glycolide)-poly(ethylene glycol) (PLGA-PEG), followed by surface functionalization with the A10 Apt, with affinity to the extracellular domain of PSMA [23]. Using this method, the hydrophilic PEG polymer protrudes outward on the surface of the polymeric NP core, and the terminal carboxy functionalities can then be conjugated with the A10 PSMA Apt targeting ligand. Additionally, the negative charge of the carboxylate groups prevents the nonspecific interaction of the Apt molecules with the surface of the NPs [20].

The efficacy of the Apt NPs was investigated using a mouse xenograft model of PCa and the NPs were injected intratumorally in LNCaP generated solid tumors (Fig. 15.2). PCa was induced in mice by implanting LNCaP prostate epithelial cells s.c. in the flanks of nude mice and allowing the tumors to develop to appropriate sizes (∼300 mm³). The tumor size and weight were monitored for up to 109 days and results revealed that a single intratumoral administration of Dtxl–NP–Apt Bioconjugate NPs was significantly effective at tumor size reduction compared to the nontargeted NP controls. The Apt conjugated NPs are believed to bind to PSMA on the surface of the LNCaP cells and become endocytosed, allowing their cytotoxic Dtxl cargo to be delivered post-endosomal release. This study demonstrated how the therapeutic index of Dtxl can be improved with the Dtxl-encapsulated PLGA-PEG-Apt NPs showing almost complete tumor reduction and 100% survival, compared with the survivability of 57% for nontargeted PLGA-PEG NPs and 14% for Dtxl alone (Fig. 15.2) [23]. The materials used in the development of these bioconjugated NPs are FDA approved and the Apts are small in size, relatively stable, nonimmunogenic, and easy to synthesize, which can facilitate the translation of these nanomedicines into clinical practice [23]. However, it is important to note that, akin to other nucleic acid-based ligands, Apts may also suffer from nonspecific interactions after systemic administration and thus it is important to engineer optimal NP surface properties and ligand densities that lead to effective targeting of nanomedicines at sites of disease.

In addition to Dtxl delivery, Apt–NP bioconjugates have also been developed for the delivery of cisplatin to PCa cells [24]. Cisplatin is the most potent member of the Pt anticancer drug family, and its use in PCa therapy is therefore highly attractive. We devised a strategy for cisplatin therapy in PCa that employed Pt chemistry and NP delivery vehicles [24]. In this study, a hydrophobic platinum (IV) prodrug was synthesized for encapsulation into polymeric PLGA-b-PEG NPs using the nanoprecipitation method, which resulted in highly loaded NPs of appropriate size. The NPs were then targeted to PSMA by decorating the surface of the particles with A10 Apts that specifically bound to the extracellular domain of PSMA. The Apt-facilitated cellular uptake of the Pt(IV)-encapsulated NPs by PSMA expressing LNCaP cells via endocytosis was demonstrated using an antibody specific for endosome formation. The Apt-derivatized Pt(IV)-encapsulated NPs were shown to be significantly superior to cisplatin or nontargeted NPs in inhibiting LNCaP cellular growth. Moreover, by encapsulating a cisplatin prodrug, PLGA-PEG-Apt NPs displayed significant dose-sparing, with equivalent antitumor efficacy in LNCaP xenografts achieved at only a third of the conventional administered dose of free cisplatin (0.3 mg/kg vs. 1 mg/kg) [25].

Newly developed nanomedicines have shown promising results in the treatment of PCa, by achieving an improvement in drug biodistribution and therapeutic index. In particular, targeted NP delivery has led to the improvement of site-specific drug delivery. In addition to exploring further specific prostate cell-surface antigens and the delivery of improved PCa drugs, other avenues such as combination therapies, precisely engineered nanomedicines, and the use of smart theranostic NPs also present exciting and promising avenues for PCa treatment and monitoring in the future.