Imaging and drug delivery using theranostic nanoparticles^☆

2.1. Optical imaging

Optical imaging is one of the more common modalities used in research. Optical imaging utilizes photons emitted from bioluminescent or fluorescent probes. It has advantages over other imaging modalities in that the detection of low energy photons is relatively inexpensive; furthermore, the spectrum from visible to near-infrared (NIR) light provides good spatial resolution, without exposure to ionizing radiation. Unfortunately, this modality suffers from poor tissue penetration (0–2 cm) and fluorescent imaging is highly susceptible to noise due to the tissue scattering of photons in the visible light region (395–600 nm) [9]. Optical imaging also suffers from significant background because of tissue autofluorescence and light absorption by proteins (257–280 nm), heme groups (max absorbance at 560 nm), and even water (above 900 nm) [11]. Despite these challenges, the NIR (700–900 nm) window has the advantages of reduced autofluorescence, reduced tissue scattering, and greater depth of penetration, which is most suitable for in vivo imaging.

2.2. Magnetic resonance imaging (MRI)

The basis for MRI signal is the precession of water hydrogen nuclei within an applied magnetic field. After application of radiofrequency pulses, the relaxation process through which the nuclei return to the original aligned state can be exploited to produce an image [12]. To enhance the differentiation between tissues, contrast agents are used to shorten the relaxation parameters (T1 and T2) of water. Paramagnetic molecules such as gadolinium (Gd) and manganese (Mn) can be tagged onto small molecules, macromolecules, or nanoparticles. Conversely, magnetic iron-oxide nanoparticles are inherently superparamagnetic and can be tagged at their surfaces with cargo. Compared to radionuclide or optical imaging, MRI has good spatial resolution; however, it suffers from low sensitivity. To offset this, relatively high concentrations of contrast agents are required to produce a detectable signal. The administration of high doses of these contrast agents leads to concerns over accumulation and toxicity, which have become a significant issue for Gd (III) complexes. While Gd (III) provides superior contrast for tumor and vascular imaging, problems with slow excretion and toxicity due to long-term accumulation may hinder its future development in the clinic.

2.3. Radionuclide-based imaging

At the other end of the electromagnetic spectrum, γ-ray emissions are the basis for SPECT and PET. For both modalities, radiopharmaceuticals are administered that can be detected by a camera. Unlike MRI, SPECT and PET images are acquired over a nominally low background signal and require little signal amplification since the gamma rays have energies in the megavolt range. In comparison to PET, SPECT is approximately ten times less sensitive; however, SPECT is advantageous because it enables concurrent imaging of multiple radionuclides. Additionally, SPECT is more widely available than PET, and SPECT radionuclides are simple to prepare and usually have a longer half life than PET radionuclides. SPECT and PET are quantitative techniques, which is an advantage over other modalities such as MRI and optical imaging [13]. Although PET suffers from poor spatial resolution, this can be overcome by hybrid imaging, where PET is used to track molecular events and high resolution CT is used to localize events. While this review specifically focuses on the diagnostic application of radionuclides, it should be noted that certain radionuclides have also been used for radiotherapy [14].

2.4. Computed tomography (CT)

While radionuclide-based imaging is adept at providing information about physiological processes, modalities like CT can provide complementary anatomical information. CT measures the absorption of X-rays as they pass through tissues. The ability of CT to distinguish tissues is based on the fact that different tissues provide distinct degrees of X-ray attenuation, where the attenuation coefficient depends on the atomic number and electron density of the tissues. Differences in absorption between bone, fat, air, and water produce high contrast images of anatomical structures [159].

Currently, CT contrast agents are typically low molecular weight and are characterized by rapid extravasation and clearance. On the other hand, macromolecular and nanoparticulate agents have exhibited superior prolonged presence in the blood pool, which may make them suitable agents for vascular CT. Nanoparticles containing electron dense elements with high atomic number such as iodine, bismuth or gold have been proposed as CT contrast agents [160]. CT contrast requires high concentrations of these elements; therefore, most research is based on solid nanoparticles or liposomes containing iodinated molecules [15]. Gold nanoparticles have gained popularity as CT contrast agents, although it is unclear if gold is an optimal material for physiological use [160,161].

2.5. Ultrasound

US is one of the most common clinical imaging modalities due to its low cost, speed, simplicity, and safety. In this modality, a transducer which emits high frequency sound waves (>20 kHz) is placed against the skin and US images are obtained based on the sound wave reflected back from the internal organs [162]. US contrast agents can improve imaging by introducing a material with different acoustic properties from that of tissues, such as gas [16].

Nanoparticulate-based contrast agents for the imaging modalities discussed are in various stages of development. Typically, the type of nanoparticle used depends on the imaging modality. The next section discusses the types of nanoparticles under development as TNPs.

3.1. Drug conjugates and complexes

In contrast to self-assembled structures like vesicles and micelles, complexation and covalent conjugation are other direct routes to prepare nanoparticles. Drug complexes rely on reversible interactions between carrier and drug, whereas drug conjugates utilize covalent interactions. Drug conjugates can be prepared using many chemical pathways, which often depends on the chemistry of the drug as well as the carrier. The two major classes of drug conjugates currently under development for theranostic use include protein and peptide-associated and polymer-associated drugs. As with vesicular structures, the effectiveness of a drug conjugate is related to its ability to improve therapeutic index relative to free drug, generally, by reducing toxicity and/or improving efficacy.

Proteins and peptides are versatile materials that can form nanoparticles in several ways, including complexation. The best example of a successful protein nanoparticle already in clinical use is Abraxane™, a formulation of paclitaxel reversibly bound to 130–150 nm albumin nanoparticles via high pressure homogenization [18,19]. Abraxane™ outperformed conventional paclitaxel at an equitoxic dose while decreasing toxicity, due to longer circulation time and decreased offtarget activity [2,18].

Synthetic polypeptides have been used in a similar way. Doxorubicin (Dox) has been chemically conjugated to chimeric polypeptide molecules using thiol chemistry, which drives nanoparticle assembly, increases plasma circulation of drug, and cellular internalization. This increased anti-tumor efficacy and reduced toxicity relative to free drug [20–23]. For drug conjugates and complexes containing Dox, the intrinsic fluorescence of the drug itself may allow optical imaging through thin tissues; however, to make imaging more quantitative it will likely be necessary to modify these conjugates with better contrast agents.

As an alternative to polypeptide polymer, many synthetic polymers have been developed as polymer-drug conjugates. One of the most widely explored polymers is N-(2-hydroxypropyl)methacrylamide (HPMA). HPMA drug conjugates have been studied for over thirty years as HPMA is non-toxic, non-immunogenic, and stable in systemic circulation [24,25]. The goals of polymer-drug conjugation include increasing circulations times, decreasing toxicity, allowing passive tumor targeting, and incorporating various functionalities, such as an active targeting moiety or contrast agent [26–28]. Many different contrast agents have been associated with HPMA copolymer conjugates for in vitro and in vivo molecular imaging [29]. An HPMA-Dox conjugate labeled with I-131 has even been evaluated in a phase 1 clinical trial [30].

Two schemes for functionalizing HPMA copolymers with diagnostic and therapeutic moieties are via copolymerization and chemical conjugation. While the latter method allows direct conjugation of these contrast agents and therapeutics to the copolymer, it suffers from low conjugation efficiency and difficulty in controlling the degree of conjugation. Even so, two chemotherapeutic agents (Dox and gemcitabine) along with I-131 have been successfully conjugated to a HPMA copolymer using this method [31]. Conversely, copolymerization may be the preferred conjugation method as it offers the flexibility to conjugate both single or multimodal diagnostic agents and chemotherapeutic drugs. An RGD targeted polymer labeled with In-111 and a HPMA construct with Gd was prepared by copolymerization [32,33].

Polymer-drug conjugation is a relatively reliable method to improve pharmacokinetics and increase therapeutic index; non-covalent complexes have a similar effect. Albumin-bound paclitaxel is routinely used in the clinic. The barriers to drug conjugate and complex development into TNPs include the necessity for efficient drug loading and minimization of complexity to control production cost and reliability. Both natural and synthetic polymers are excellent platforms for overcoming these barriers.

3.2. Dendrimers

Polymeric dendrimers are hyperbranched nanostructures that can be controlled in size by controlling the number of polymerization generations. As polymerization progresses, a small, planar molecule transforms into a spherical nanostructure, with cavities where therapeutics and contrast agents can be grafted with great loading efficiency [34]. Dendrimer polymerization and synthesis can be regulated to precisely control the molecular weight and chemical composition of the final product [35]. From the perspective of controlling polymer structure and polydispersity, dendrimers may be ideal platforms for TNPs.

The only dendrimer-based formulation to enter clinical trials thus far is VivaGel™, which uses the dendrimer as a therapeutic agent rather than a carrier — it is currently being evaluated for safety and efficacy as a microbicide [36]. In a recent preclinical study, a fifth generation polyamido-amine (PAMAM) dendrimer conjugated to fluorescein isothiocyanate for imaging and recombinant Fibroblast Growth Factor-1 for tumor targeting was developed to be used to track targeting and cellular internalization [37]. The same group demonstrated a novel release mechanism from a Dox-loaded, folate targeted dendrimer, whereby Dox is conjugated via a photo-sensitive linker, allowing light-controlled drug release [38]. Another group has demonstrated the in vivo and in vitro efficacy of biodegradable dendrimers, with Dox conjugated via pH-sensitive hydrazone linkers, which showed good biocompatibility, biodistribution, and pharmacokinetics [39–41].

Additions of Gd or iron-oxide are necessary to create a dendrimer-based MRI contrast agent. Wiener and coworkers were the first to demonstrate the viability of dendrimers as MRI contrast agents [42]. The presence of multiple functionalization sites on dendrimers allows attachment of targeting ligands like monoclonal antibodies, peptides and folate [43]. Konda and coworkers, synthesized a fourth generation PAMAM dendrimer with Gd-DTPA chelates. These dendrimers were modified with folic acid to target up-regulated receptors in ovarian cancer [44]. Targeted dendrimers produced site-specific enhanced image contrast and drug localization, in another formulation [45]. To overcome toxicity with macromolecular Gd (III) complexes, Xu and coworkers synthesized a dendrimer with a biodegradable spacer that rapidly eliminates the contrast agent [46]. High relaxivity, efficient loading, and enhancement of tumor and blood pool imaging were observed. Despite encouraging results, the authors concluded that the conjugate is not suitable for further development due to the inherent toxicity of PAMAM dendrimers. Instead, they recommend developing non-toxic, biodegradable dendrimers as a safer alternative which are under development by Szoka and coworkers [35].

The starburst structure of dendrimers allows multivalent attachment of chelators and targeting moieties. Almutairi and coworkers demonstrated highly selective imaging with Br-76 labeled dendritic nanoprobes in an angiogenic mouse model [47]. Attempts have also been made by Kobayashi and coworkers to target delivery of In-111 generation-4 PAMAM dendrimer in a mouse tumor model [48]. Biodistribution indicated rapid blood clearance and preferential accumulation at the tumor site. A similar study by the same group using a second generation PAMAM dendrimer with Y-88 yielded similar results [49]. Yardanov and coworkers characterized an iodinated dendritic nanoparticle, G4-(DMAA-IPA)₃₇ as a CT imaging agent [50]. The construct consists of DMAA-IPA covalently conjugated to a fourth generation PAMAM dendrimer. Due to the multiplicity of DMAA-IPA conjugation, high capacity loading of iodine was achieved, which enhanced contrast for CT.

Dendrimers have been in development for years and have proven to be successful drug and imaging agent carriers in a number of preclinical studies, including, tumor regression, gene delivery, and molecular imaging [6,34,51]. Barriers to address prior to theranostic use of dendrimers include toxicity of the nanoparticle, its polymeric component, as well as the off-target effects of the contrast and therapeutic agent.

3.3. Vesicles

Vesicular structures have been extensively studied as clinical therapeutics and experimentally as diagnostic nanoparticles. The two major classes of vesicles, liposomes and polymer vesicles, have the capacity for covalent and non-covalent encapsulation of both hydrophobic and hydrophilic cargo. Thus, vesicles are well suited to co-deliver therapeutic and diagnostic agents, which can differ in their physicochemico properties. Liposome formulations for therapeutic and diagnostic delivery have been in development for almost half a century, and have produced clinically successful applications [52]. Polymeric vesicle structures, or polymersomes, have been described more recently, which have unique potential as drug and contrast agent carriers [53].

The best example of a clinical vesicle formulation is Doxil™, also known as liposomal Dox. Encapsulation of Dox in liposomes shielded by polyethylene glycol (PEG) prolongs systemic drug circulation, amplifies safety, and improves the therapeutic index relative to free Dox [6,19]. Moreover, liposomes that co-encapsulate multiple agents are undergoing clinical evaluation. CPX-351, a liposomal formulation with a synergistic molar ratio of 5:1 of cytarabine and daunorubicin, respectively, is currently in Phase II clinical trials for the treatment of acute myeloid leukemia [54]. These liposomes are prepared with cytarabine in the lipid mixture and loaded with daunorubicin after formation, such that the desired loading efficiencies and molar ratios are achieved [55]. Like Doxil™, CPX-351 has prolonged pharmacokinetics, leukemia-selective uptake and cytotoxicity, and improved in vivo efficacy, possibly explained by greater drug exposure in the bone marrow and plasma [56,57]. Although CPX-351 is a purely therapeutic formulation, it has been fluorescently labeled and even imaged using the intrinsic fluorescence of the anthracycline, daunorubicin — allowing the possibility of in vivo imaging [54,56].

There are multiple examples of liposomal formulations that incorporate targeting, therapeutic, and imaging functionalities [58–62]. As diagnostic agents, liposomes have been labeled with quantum dots (QD) Mn and Gd (both aqueous [63] and chelated [64]), radionuclides such as Ga-67, In-111 and Tc-99 m (anchored to either the lipid bilayer or within the aqueous core), iodine-based agents [15], or even gas (echogenic liposomes [65,66]). As a platform, liposomes can incorporate multiple functionalities without significantly compromising their stability in the body, which makes them suitable nanoparticles for theranostics.

While liposomes are composed from phospholipids chemically similar to those in eukaryotes, some of their most significant enhancements result from polymeric modification. In fact, vesicles can arise from the self-assembly of bilayers composed exclusively from non-lipid polymers. Polymer vesicles or polymersomes have been formed using diblock or triblock copolymers of various polymeric materials including polylactic acid (PLA), polyglycolic acid (PGA), polylactic-co-glycolic acid (PLGA), polycaprolactone (PCL), chitosan, and PEG [67–70]. The principle of these block copolymers is that the hydrophobic block self-associates to form a core membrane, which is stabilized in aqueous media by the hydrophilic polymeric block. As such, polymersomes can encapsulate both hydrophobic and hydrophilic molecules with acceptable encapsulation efficiencies [71–75]. Chemotherapeutic polymersomes carrying Dox and docetaxel have demonstrated improved in vitro and in vivo efficiency, mediated by controlled release, cellular internalization, and prolonged pharmacokinetics [72,76]. A study demonstrated the specificity of an actively targeted polymeric vesicle via encapsulation and targeted release of a hydrophobic dye [77]. In addition, gene delivery has been demonstrated in vitro and in vivo, while simultaneously allowing molecular imaging of the polymersome [74,78,79]. Gd-DTPA and iron-oxide loaded PLGA nanoparticles have been used with US and MRI [80,81]. Recently, a construct with similar efficacy and rapid clearance was reported to limit toxicity [82]. Here, superparamagnetic iron-oxide (SPIO) was formulated in nanoparticles consisting of polylactic acid and d-α-tocopherol PEG 1000 succinate. Sequestering the SPIO within this nanoparticle helps to overcome undesirable liver accumulation, which potentially may reduce hepatotoxicity.

Both liposomes and polymersomes are versatile and tunable vehicles for drug delivery and imaging. Many strategies used to form and load liposomes can also be applied to polymersomes. Furthermore, polymersomes offer enhanced control over the membrane thickness and stability, which may offer pharmaceutical advantages. When incorporating new materials and chemistries on these complex particles, it is important to consider their limitations, such as their inherent production complexity and the vehicle toxicity. Vesicular structures can efficiently load drugs and imaging agents, via either covalent or non-covalent forces, incorporate targeting strategies, prolong systemic circulation, and increase tumor accumulation.

3.4. Micelles

Micellar nanoparticles are attractive structures for carriers of drug and contrast agents because they can form relatively uniform size structures, be prepared from a variety of amphiphilic materials, increase solubility of hydrophobic molecules, and incorporate multiple functionalities into a single structure. Lacking an aqueous core, the drug and contrast agent must be bound to the polymer prior to formation, conjugated to an anchor molecule, or entrapped and associated within the dense, hydrophobic core of the micelle [83]. Micellar structures, including polymeric micelles, have been extensively studied as drug carriers [84].

Like polymersomes or polymer-drug conjugates — similar polymeric materials can be used to form polymeric micelles. In general, a hydrophobic block of the copolymer forms the micellar core, while a hydrophilic portion forms the corona. This corona (commonly consisting of PEG, HPMA, or equivalent hydrophilic polymer) confers these micelles with biocompatibility, stealth-like properties, and a platform for functionalization [85–87]. Genexol-PM™ is a formulation of paclitaxel encapsulated in a polymeric micelle formed by the solid dispersion technique from the biodegradable block copolymer, monomethoxy poly(ethylene glycol)-block-poly(d,l-lactide) [88]. Several clinical trials have already been performed evaluating and validating the safety and efficacy of Genexol-PM™ in metastatic breast cancer, solid tumors, and non-small-cell lung cancer [89–91].

Similarly, contrast-loaded micelles can be used for visualizing numerous organs, tissues and disease sites. Membranotropic chelating agents such as DTPA stearylamine (DTPA-SA) [92] or DTPA-phosphatidylethanolamine (DTPA-PE) [93] have been developed whereby the lipid part of the molecule can be anchored in the micelle’s hydrophobic core while a more hydrophilic chelate is localized on the hydrophilic corona. Similarly, polychelating amphiphilic polymers (PAP) that possess multiple side groups for chelation can be anchored to the micellar surface. Different combinations of chelators and hydrophobic anchors have been tested in the preparation of In-111, Tc-99 m and Gd liposomes [94–96]. Chelates with high stability under physiological conditions are preferred so as to avoid toxicity [97–99]. Due to multiplicity of attachment, PAPs can improve image quality at a low micellar dose. Trubetskoy and coworkers incorporated amphiphilic chelating probes (In-111-DTPA-PE and In-111-DTPA-SA) into PEG-PE micelles of 20 nm diameter and used these micelles for γ-scintigraphy of the lymphatic system after subcutaneous administration in rabbits [100]. Moreover, PEG-PE-DTPA micelles labeled with In-111 and an antibody have been used to image a murine model of lung carcinoma [101]. Additionally Trubetskoy and coworkers have also developed iodine-containing micelles for use with CT. Methoxy-polyethylene glycol and tri-iodobenzoic polylysine micelles improved X-ray signal when in rats and rabbits [102].

Due to ease of formation, stability, ability to encapsulate hydrophobic molecules, and therapeutic success in preclinical and clinical studies, polymeric micelles are widely accepted as viable drug and imaging agent delivery systems [86,103–107]. However, the stability of micelles depends on their critical micelle concentration, which makes them prone to exchange their components with other physiological membranes.

3.5. Core–shell

Core–shell structured nanoparticles include particles made with a wide variety of materials including QDs, metals, metal oxides, and polymers among other materials. Their structure resembles a micelle in that there is a core surrounded by a hydrophilic shell. However, most of the core–shell particles relevant to TNPs are stabilized by covalent or ionic bonds, and not hydrophobic effects. There have been advances in the chemical modification of gold and other materials to increase their physiological stability and control drug association and release [17,108]. Core–shell nanoparticles offer stability, modifiable surface properties, and can even produce contrast depending on their composition, size, and shape [109,110].

Metal nanoparticles are attractive options for drug and contrast agent delivery because they are a stable platform on which multiple functionalities can be grafted, some can be imaged directly, and certain formulations allow magnetic-directed guidance [110,111]. The latter can be used to localize drugs, genes, and contrast agents to a specific location. Various SPIO-based nanoparticles have been synthesized that are useful for magnetic guidance, can be imaged, and have surface chemistry that can be modified in order to achieve acceptable biodistributions, while limiting clearance by the reticuloendothelial system (RES) [111]. Also, the superparamagnetism of the iron-oxide nanoparticle allows direct imaging with MRI. SPIO have been widely used to provide contrast for T2 weighted MRI. SPIO can be used alone or attached to other nanostructures.

In a recent study, to demonstrate the viability of these nanoparticles, paclitaxel and a dye were co-encapsulated in the polymer matrix that surrounds the SPIO core. These particles showed dose-dependent internalization and cytotoxicity via multimodal imaging of the dye and SPIO [112]. There are even FDA-approved SPIO nanoparticles, Feraheme™ (Ferumoxytol) is a SPIO nanoparticle with a carbohydrate coating. Feraheme™ can be used for intravascular imaging as well as treating iron deficiency anemia [113,114]. A similar vascular application was reported by Senpan et al. [115]. SPIO can also be manipulated to label specific cell types, for monitoring their in vivo behavior [116].

Fluorescent semiconductor QD are nanocrystals consisting of atoms such as cadmium (Cd), zinc (Zn), selenium (Se), telerium (Te), and sulfur (S). However, the most commonly studied QD formulations in biological applications contain a cadmium selenide core and a zinc sulfide shell, which is surrounded by a hydrophilic polymer that can be used to load therapeutics and conjugate targeting ligands [117–119]. QDs can be tuned to emit at between 450 nm to 850 nm by changing their size or chemical composition. QDs improve signal brightness up to 10–100 times when compared to organic fluorophores. Additionally, they have a wide excitation and narrow emission window, fluoresce with high efficiency, and resist photobleaching. Thus far, QDs have also been applied to cancer research in animals, including sentinel node mapping and in vivo molecular imaging of tumors [120–124]. QDs have been developed for single molecule labeling, to observe biomolecular dynamics in live cells [125]. The main drawback of QDs is their toxicity. Oxidative degradation of its heavy metal core will lead to the release of metal ions, which bind to sulfyhydryl groups on intracellular proteins and disrupt the function of subcellular organelles.

AuroLase™ is a gold nanoshell formulation in clinical trials for photothermal tumor ablation [126]. The formulation is composed of a gold nanoshell grown around a colloidal silica particle, with thiolated-PEG assembled on the exterior nanoshell surface [127]. After intravenous administration, these nanoparticles passively accumulate in tumor and can be monitored via MRI. After the particles have localized, the tumor is exposed to NIR light, which causes thermal ablation [127]. The in vivo imaging and therapeutic effect of this nanoparticle system has been evaluated in a canine and murine tumor models, and is entering the clinic [126–128].

New synthetic approaches have yielded highly luminescent gold nanoparticles with emissions ranging from 400 to 1200 nm [129,130]. Nanoparticles with this NIR emission makes them suitable for optical imaging as they are less affected by autofluorescence and provide better contrast of the target tissues. Similar to QDs, the emission spectra of gold nanoparticles are size dependent. In addition to optical contrast, gold nanoparticles have been used as CT contrast agents in small animals with little toxicity [131]. Gold is a better CT contrast agent than iodine because gold has a higher atomic number and electron density, which leads to stronger X-ray attenuation and greater contrast. Furthermore, imaging gold at 80–100 keV reduces interference from bone and tissue, which would reduce the radiation exposure to patients. Experiments performed in mice using gold nanoparticles (1.9 nm, ~50 kDa) exhibited good stability, high X-ray absorption, low toxicity, prolonged plasma half life, and enhanced CT contrast of the vasculature, kidneys, and tumor in a mouse model [131].

These core–shell structured nanoparticles make for a very robust therapeutic and contrast agent delivery vehicle due to the vast array of elements that can be used as the core and shell components. Despite the advantages already mentioned, these core–shell nanoparticles suffer from some inherent shortcomings such as toxicity, rapid clearance, and poor biodistribution and biodegradation. Various strategies are in development to circumvent these problems [80,81].

3.6. Microbubbles

Microbubbles and nanobubbles are spherical cavities filled with gas and are usually <10 μM in size. These microbubbles are of varying chemical composition and can be induced to expand and contract (resonate) in the presence of US delivered at the resonance frequency of the microbubbles. The most common gasses used for this purpose in commercially produced contrast agents include perfluorocarbons (PF), such as octofluoropropane (C₃F₈), decafluorobutane (C₄F₁₀) and sulfur hexafluoride (SF₆). The shell encapsulating the gas can be composed of denatured albumin, phospholipids, or various polymers — this shell can then be functionalized to associate with targeting molecules and therapeutics [132,133]. Microbubbles are used as contrast agents for imaging inflammation, angiogenesis, intravascular thrombus and tumors. More recently, microbubbles and nanobubbles are being investigated as potential drug carriers in addition to their role as contrast agents [134].

A study conducted by Gao and coworkers demonstrated the incorporation of Dox into the walls of a polymeric microbubble. PEG-PLLA or PEG-PCL block copolymers were used to form micelles that encapsulate Dox. Then perfluoropentane (PFP) is added and the solution is sonicated — resulting in a mixture of Dox-loaded micelles and Dox-loaded, PFP-encapsulating nanobubbles [135]. As demonstrated in vitro and in vivo using a murine tumor xenograft model, this Dox-loaded nanobubble/micelle mixture was shown to passively accumulate in tumor tissue and coalesce to form microbubbles, which cavitate and collapse upon tumor-directed US, resulting in drug release [135,136]. This selective accumulation and drug release provided effective Dox delivery and cellular uptake, while simultaneously allowing for molecular imaging of the nano/microbubbles. Another study demonstrates the application of microbubbles for theranostics by using targeted gold nanoparticles to create plasmonic nanobubbles in situ after a short laser pulse, which provides contrast for US imaging (diagnostic). This strategy disrupts cellular membranes due to rapid expansion to microbubble size and subsequent collapse upon more laser exposure (therapeutic) [137].

Microbubbles and nanobubbles have been developed as contrast agents for US imaging for a number of years, but their therapeutic potential has recently begun to mature. The use of these nanoparticles as therapeutic agents is still relatively new and requires further evaluation in more complex systems.

3.7. Carbon nanotubes

Carbon nanotubes (CNTs) are cylindrical tubes composed solely of carbon and can either be formed single-walled or multi-walled, for more stability [138]. These CNTs are being investigated as TNPs because of their tunable properties and ability to incorporate multiple functionalities.

Amine-functionalized single-walled CNTs were conjugated via amide linkages to a cisplatin-folic acid derivative as targeted therapeutics, which outperformed a conventional cisplatin formulation [139]. Another study used carboxylated single-walled CNTs to conjugate cisplatin and epidermal growth factor to target cells overexpressing the epidermal growth factor receptor. The targeted CNT-cisplatin conjugate improved the selective killing of the targeted tumor cells [140]. Many studies have confirmed CNTs to be effective carriers of therapeutics and imaging agent in vitro, however, in vivo studies have lead to concerns regarding the associated toxicities [141–143].

As with QDs and gold nanoparticles, a broad excitation profile and high absorption coefficient can be obtained by tuning CNT over a wide range of wavelengths. Relatively few studies have been conducted utilizing CNTs in optical imaging. However, NIR and Raman signals from CNTs have been detected from within subcutaneous tumors in mice [144]. CNTs provide poor MRI contrast, but they can be functionalized to be made detectable. Richard and coworkers reported functionalization of CNT with amphiphilic Gd (III) chelates and studied their effect on water proton relaxation in vitro and in vivo [145]. Functionalization of a single-walled CNT surface allows conjugation of In-111 labeled DOTA chelate an anti-CD20 monoclonal antibody for targeted imaging of lymphoid malignancies [146]. A similar study conducted by Liu and coworkers demonstrated that radiolabeled single-walled CNTs can be detected by PET scanning [147]. They also showed that single-walled CNTs conjugated with RGD peptide were able to target and image α_vβ₃ integrin expression in tumors.

CNTs have demonstrated preclinical efficacy for therapeutic and contrast agent delivery — making them interesting nanoparticle formulations for theranostic use. However, the primary barriers that must be overcome include route of biodegradation and reduction of off-target toxicity.