Image guided therapy: The advent of theranostic agents

Journal of Controlled Release 161 (2012) 328–337

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Review

Image guided therapy: The advent of theranostic agents Enzo Terreno a,⁎, Fulvio Uggeri b, Silvio Aime a a b

Department of Chemistry and Molecular & Preclinical Imaging Centers, University of Turin, Italy Bracco Imaging S.p.A., Centro Ricerche, Colleretto Giacosa (TO), Italy

a r t i c l e

i n f o

Article history: Received 9 February 2012 Accepted 15 May 2012 Available online 22 May 2012 Keywords: In vivo imaging Drug delivery and release Image-guided therapies Nanomedicine

a b s t r a c t Theranostic agents represent a recently introduced class of imaging probes designed to offer to pharmacologists and physicians a robust tool for minimally invasive in vivo visualization of drug delivery/release and therapy monitoring. By means of these agents, novel strategies able to integrate diagnosis and therapy could be developed. This highly interdisciplinary research field is one of the more innovative products resulting from the synergism between molecular imaging and nanomedicine. Potential applications of theranosis include the in vivo assessment of drug biodistribution and accumulation at the target site, visualization of the drug release from a given nanocarrier, and real-time monitoring of the therapeutic outcome. The expected end-point of theranostic agents is to provide a fundamental support for the optimization of innovative diagnostic and therapeutic strategies that could contribute to emerging concepts in the field of the “personalized medicine”. This perspective paper aims at providing the reader the basic principles of theranosis with a particular emphasis to the design of theranostic agents. © 2012 Elsevier B.V. All rights reserved.

Contents 1. Introduction . . . . . . . . . . . . . . . . . 2. Imaging guided drug-delivery . . . . . . . . . 3. Imaging drug-release . . . . . . . . . . . . . 4. Monitoring therapy by imaging . . . . . . . . 5. Theranostic agents for radiation-based therapies 6. Imaging-guided surgery . . . . . . . . . . . 7. Concluding remarks . . . . . . . . . . . . . Acknowledgments . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . .

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1. Introduction The search for novel drug delivery technologies is essentially driven by the observation that many therapeutic agents failed due to their limited ability to reach the target tissue and/or because of their poor selectivity against diseased cells. In addition, drug delivery systems can solve problems associated with drug instability in the biological environment as well as issues related to the modulation of drug

⁎ Corresponding author at: Department of Chemistry and Molecular & Preclinical Imaging Centers, University of Turin, Via Nizza, 52, 10126, Italy. Tel.: +39 0116706452; fax: +39 0116706487. E-mail address: [email protected] (E. Terreno). 0168-3659/$ – see front matter © 2012 Elsevier B.V. All rights reserved. doi:10.1016/j.jconrel.2012.05.028

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clearance and metabolism. Improved delivery systems are most in need for the treatment of cancer, respiratory and central nervous system diseases and for cardiovascular disorders. A wide variety of carriers have been investigated so far, including lipid-based self-assembled systems [1], polymeric/inorganic particles [2,3], “host–guest” supramolecular adducts [4], and naturally-occurring systems like lipoproteins [5], proteins [6], viral capsids, bacteria, and even whole cells [7]. Drugs can be released from the carrier spontaneously or through specific chemical or physical triggering stimuli like pH, enzymes, heat, ultrasound or magnetic forces [8]. Delivery/release steps can be visualized in vivo by ‘doping’ the nanomedicine with suitable imaging probes that can generate a response detectable by the currently available imaging technologies.

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This task relies on the great advances achieved in the last decade in the field of molecular imaging that has allowed the design of probes able to provide an imaging response associated with functional/ biochemical processes occurring in the vascular, extracellular and intracellular compartments [9]. The successful combination of molecular imaging and nanomedicine approaches has given a fundamental contribution to the birth of a new, highly interdisciplinary research field called theranosis [10]. Indeed the main task of the theranostic field deals with the development of innovative strategies to provide imaging tools for the in vivo visualization of drug delivery, (triggered) drug release, and therapy monitoring. An important inference of this concept is represented by the role of imaging in interventional procedures that, at present, appears to be one of the applications that is more close to clinical translation. The development of theranosis has been made possible also by the great progress in the field of imaging technologies. These achievements allow a very fast image acquisition and processing still maintaining very high levels of sensitivity and spatial resolution. Moreover, hybrid systems such as HIFU-MRI (HIFU: High Intensity Focused Ultrasound) [11], already introduced in the clinical settings for imaging-guided hyperthermia treatments, have good chances to play a relevant role in the future development of innovative theranostic agents. This survey will report on the different application areas of theranostic agents (drug-delivery, drug release, monitoring therapy, agents for radiation-based treatments, interventional imaging) with the aim of providing the reader with an overview (though non-exhaustive) of the stateof-the-art in the field, giving some insights on the future perspectives. Despite the youth age of theranosis, a number of publications have already appeared. However, most of them deals essentially with the evaluation of the imaging properties of the nanocarrier-based systems leaving the therapeutic function of the potential theranostic agent unchecked. In the present survey we have selected examples of theranostics “at work”, i.e. systems that fully exploit the synergy between the diagnostic and the therapeutic companions. 2. Imaging guided drug-delivery As anticipated above, the advent of nanomedicine was strongly driven by the need of improving drug delivery at the target site to optimize its therapeutic index. Typically, a preclinical biodistribution study leads to the quantification of the drug over time in major tissues/organs. Traditionally, these studies were carried out ex vivo through time-consuming procedures that required the sacrifice of a large number of animals. Nowadays, this can be done in vivo and in real-time using imaging methodologies and tailored imaging probes. In a typical nanomedicine approach, the nanocarrier is loaded with the drug and the imaging reporter that can be detected by the corresponding imaging technique. Due to the outstanding detection sensitivity and good quantification properties, radioisotopes that have been widely used for exvivo biodistribution studies, may have a good potential for being successfully used also in vivo. Among the many literature reports, it is worth of note the work of Li and co-workers who recently proposed a very innovative approach [12]. They genetically engineered a T7 bacteriophage and decorated it with the RGD peptide (Fig. 1, top) that is known to act as specific targeting vector toward αvβ3 integrin. The external surface of the phage was also functionalized with 64Cu tracers (based on DOTA or sarcophagine coordination cages) for PET imaging. The interest to use phages for imaging drug-delivery relies on the easy replacement of the nuclear material with a drug. Bottom of Fig. 1 reports a series of PET images obtained over time after i.v. injection of the functionalized phage in a mouse model of human glioma (U87-MG). The graphs reported on the right show the in vivo distribution of the agent (square labels) in selected organs. In a successive experiment, the injection of the PET responsive agent

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was preceded by the administration of the construct not functionalized with the PET tracer to target αvβ3 integrin without generating a PET signal. The new biodistribution data (triangle labels) were very similar in all the organs except the tumor, where a significantly lower signal was detected due to the extensive blockage of the receptors operated by the PET-silent agent. Taken together, these results outline the positive role of the targeting vector for guiding the agent to the tumor lesion and emphasize the potential of PET imaging for the quantification of the drug-delivery. Besides nuclear imaging, also other techniques strongly contributed to the growth of the theranostic field. In this respect, Magnetic Resonance Imaging (MRI) has several strong points like the outstanding spatial resolution (down to tens of μm), minimal invasiveness and excellent potential for longitudinal studies. However, MRI suffers for the relatively low sensitivity of its contrast agents, especially when the injected probe is intended to report about biological processes occurring at cellular level. A useful approach to overcome the sensitivity issue is to amplify the contrast by accumulating a large number of imaging agents at the target of interest. Therefore, nanotechnology has been under intense scrutiny for the design of MRI-based molecular imaging protocols based on either paramagnetic or superparamagnetic agents [13]. In case of paramagnetic agents (Gd(III) or Mn(II) complexes), the tight binding to the structural components of the nanocarrier results in an elongation of the rotational tumbling of the metal complex that has an additional, remarkable positive effect on the enhancement of MR contrast (especially at the clinically relevant field strengths). Most of the work reported so far in the field consisted of conjugating paramagnetic agents to soft or hard particles. The typical approach in the case of lipid-based nanocarriers (which are the most diffuse and, potentially, more close to the clinical translation) involves the use of amphiphilic complexes that can be easily loaded into micelles or liposomes [14]. Recently, our group published two examples of this class of theranostic agents. In both cases, the paramagnetic complex (GdDOTAMA(C18)2, see Fig. 2) was incorporated into the bilayered membrane of pegylated liposomes encapsulating the anticancer drugs Doxorubicin [15] or Prednisolone phosphate [16], respectively. The former study aimed at using MRI to monitor the tumor accumulation of Doxorubicin-containing stealth liposomes tagged with the Gd-based agent and exposing on their surface a tetravalent peptide able to target NCAM (Neural Cell Adhesion Molecule) receptors. Such an epitope is overexpressed in neoformed endothelium in the tumor lesion as well as in several solid tumors (e.g. Kaposi's sarcoma). The theranostic nanomedicine is schematically represented in Fig. 2. An excellent correlation between the MRI response (longitudinal water protons relaxation enhancement) and the amount of Doxorubicin

Fig. 1. Top: synthetic scheme for preparing the 64Cu-tagged phage targeting αvβ3 integrin. Bottom left: coronal μPET images of U87MG tumor-bearing (white arrow) nude mice after injection of the theranostic agent. Bottom right: biodistribution of the PET nanotracer. Squares refer to the study carried out after administration of 64Cu-loaded phage only, whereas circles show the biodistribution of the agent in mice pre-treated with targeted phages not functionalized with 64Cu. Adapted from [12].

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Fig. 2. Schematic representation of the nanomedicine used in Ref. [15]. Liposomes consisted of: i) a drug (doxorubicin) entrapped in the inner core of the vesicle as insoluble salt, ii) an amphiphilic Gd-based MRI agent incorporated in the liposome bilayer, iii) PEG chains for endowing the nanoparticles with stealthiness properties toward mononuclear phagocytes system, and iv) a tetravalent peptide (named C3d) conjugated at the terminal end of PEG.

internalized by tumor cells has been observed in vitro experiments (Fig. 3). In vivo experiments, carried out on immuno-deficient mice bearing xenografted Kaposi's sarcoma, showed a superior therapeutic effect for the targeted liposomes (evaluated by monitoring the tumor growth by MRI). A complete set of ex-vivo analysis (electron and fluorescence confocal microscopy and histology) confirmed the higher cytotoxic effect produced by the targeted nanomedicine in comparison with similar, but non-targeted, liposome used as control. However, quite surprisingly, the MRI positive contrast generated by the targeted nanomedicine was lower than that one detected for the control, Gd-loaded but untargeted, nanoparticles (Fig. 4). The observed behavior can be accounted in terms of the different distribution of targeted and untargeted liposomes in the tumor region. In fact, ex-vivo electron micrographs emphasized a predominant intracellular localization of the targeted liposomes, whereas the nontargeted nanoparticles were primarily found in the extracellular compartment. Since the contrast response decreases with the intracellular localization of the agent (more barriers to overcome for the tissutal water coupled to “quenching” effects due to the endosomal accumulation of the Gd-containing agents), the observed in vivo results can be accounted in terms of the differences in the distribution of the two systems. As one would have preferred to detect an enhanced signal for the targeted agent, this finding appears detrimental for the quantification of the delivered drug, but at the same time, it provides interesting insights on the great potential of MRI to investigate the tissue localization of the nanomedicine. Though most of the nanocarriers developed for drug delivery appears more amenable for tagging with T1 MRI agents, also iron oxide particles have been considered in the design of theranostic probes, for the visualization of therapeutic treatment in cancer models [17].

Fig. 3. Left: in vitro intracellular concentration of doxorubicin measured in Kaposi's and TEC cells after 3 h incubation with targeted liposomes (C3d), non-targeted control (PEG lipo), and doxorubicin alone. The drug dose was the same in each experiment (50 μg/mL). Right: correlation between observed longitudinal water protons relaxation rates (R1obs—s−1, measured at 7 T) and amount of doxorubicin delivered to target cells by targeted liposomes after 3 h of incubation at increasing concentrations of Gd (0.01–0.1 mM). Adapted from Ref. [15].

Fig. 4. Top: T1-weighted MR images of Kaposi's sarcomas (circled in red) acquired 24 h post administration of targeted liposomes (C3d) and control nanoparticles (PEG lipo). Bottom left: MR signal enhancement measured in the tumor region. Bottom right: quantification of doxorubicin from tumor tissue 24 h post administration of the nanomedicine. Adapted from Ref [15].

Recently, Chourpa and colleagues prepared superparamagnetic iron oxide (SPIO) particles covalently loaded with Doxorubicin and Folic acid [18]. Though the cytotoxicity of the construct was not tested, tumor cells took up the targeted theranostic agent much more avidly than the control non-targeted agent. Despite the reduced penetration of exciting and emitted photons, which limits the clinical translation of the technique, also Optical imaging has been widely used to assess the efficacy of theranostic agents. An interesting example that illustrates the potential of this approach at preclinical level has been recently reported by the Weissleder's group [19], who synthesized a multimodal theranostic nanoagent (CLIOTHPC) for the treatment of inflammatory atherosclerosis based upon magneto-fluorescent nanoparticles. SPIO nanoparticles were conjugated with a near-infrared (NIR) fluorophore (AlexaFluor 750) and a light activated therapeutic (chlorin derivative), which allow for both the optical determination of agent localization and phototoxic activation. The resulting agent was readily phagocytized by murine macrophages in vitro, and resulted to be highly phototoxic. The in vitro results were confirmed in vivo, where the theranostic agent localized within macrophage-rich atherosclerotic lesions that were imaged by intravital fluorescence microscopy (Fig. 5). The exposure of the lesion to light photons activated the therapeutic response that, in turn, led to the stabilization of the lesion due to the eradication of inflammatory macrophages. 3. Imaging drug-release The therapeutic efficacy of a drug transported by a nanocarrier is not uniquely dependent on the accumulation at the pathological target, but it can be strongly affected by the release property of the

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Fig. 6. Left: schematic representation of temperature-induced release from a TSL containing doxorubicin (red dots) and Gadoteridol (yellow dots). Right: morphological MR images of tumor-bearing rats (2 treated and 1 control, upper row) and T1 maps of the tumor and leg overlaid on anatomical images taken at the time points reported on the left of the images. One HIFU-treated animal (rat 1) showed a large T1 response in the tumor, while another animal showed a much less response (rat 2). Adapted from [24].

Fig. 5. In vivo intravital fluorescence microscopy of the theranostic nanoagent CLIO-THPC to carotid atheroma. Top: emission of the NIR dye AF750 demonstrating the particle uptake by the plaque. Middle: fluorescence angiogram utilizing standard fluorescein-labeled dextran for outlining vasculature. Bottom: merged images. Adapted from [19].

whole construct. Therefore, in addition to set up imaging protocols for the visualization of drug-delivery, it is of interest to develop imaging strategies to get information on the step associated to the drug release. This task is very challenging and it requires the design of ‘smart’ imaging probes able to generate contrast changes in response to the release of the drug from the nanocarrier. Again, whereas the release process can be quite easily monitored in vitro, the challenge posed by theranostics is to replicate the behavior in vivo. Nuclear imaging technologies (PET, SPECT) as well as CT are not suitable to tackle this task because the imaging response generated by radioactive- or X-ray opaque tracers is substantially independent from the carrier-bound/carrier-free condition. On the other hand, the great sensitivity to the microenvironment shown by MR and Optical imaging probes makes these modalities particularly suitable to assess the details of the drug release process. An example is represented by the use of MRI to visualize the release of drugs from nanovesicular carriers like liposomes, which are by far the most used nanocarriers for drug-delivery purposes. The encapsulation of a hydrophilic paramagnetic MRI agent in the aqueous core of the nanoparticle may significantly reduce the ability of the probe to generate contrast [20]. This behavior is explained by taking into consideration that the contrast efficacy of such chemicals is proportional to the ability of bulk solvent water molecules to ‘feel’ the presence of the unpaired electrons of the paramagnetic metal ion. On this basis, the presence of membranes/barriers that reduce the accessibility of bulk water to the metal ion is detrimental. Hence, the encapsulation of the probe in liposomes with a low water permeability (which are those typically used in clinical applications [21]) provides an intriguing opportunity to observe a contrast enhancement when the probe (and the associated drug) is released from the carrier. This approach was smartly tested by Dewhirst and colleagues, who modified the typical procedure for remotely loading Doxorubicin in thermosensitive liposomes (TSLs) replacing ammonium ion with paramagnetic Mn(II) ion [22]. The release of the drug was locally triggered by heating and it was accompanied by the detection of a good contrast enhancement that was higher than that observed using temperature insensitive liposomes as control. These observations were corroborated in a follow-up study, in which the MRI response

nicely correlated with Doxorubicin tumor content. This seminal paper was followed by more recent investigations where Doxorubicincontaining TSLs was co-loaded with the clinically approved Gd(III) agent Gadoteridol [23]. Researchers at Philips provided an elegant in vivo proof-of-concept to demonstrate the ability of this theranostic agent visualize HIFU-mediated drug release [24]. HIFU trigger was applied for 30 min in 9 L rat tumors using a clinical MRI-HIFU system. The local release of Gadoteridol was monitored with interleaved T1 mapping of the tumor tissue (Fig. 6). A good correlation between the imaging response and the tumor uptake of Doxorubicin and gadolinium was found, thus demonstrating once again the potential of MRI for in vivo monitoring of drug release. Moreover, the reported data emphasized the large intra-animal variability for which the tumor uptake of the nanomedicine (and the corresponding imaging outcome) can significantly differ from animal to animal, thereby envisioning interesting insights for the stratification between responding and non-responding individuals. MRI is a highly versatile modality and it offers alternative ways to T1 contrast to report about drug release. For instance, other NMR active nuclei and other contrast mechanisms can be used. An example was reported by Langereis et al., who designed a TSL whose inner core was loaded with a paramagnetic shift reagent (SR), able to significantly affect the resonance frequency of the intraliposomal water protons, and with a 19F-containing molecule (Fig. 7) [25]. This dual theranostic agent can be visualized by applying a CEST (Chemical Exchange Saturation Transfer) imaging procedure at physiological temperature (i.e. below the transition temperature that causes the release of the encapsulated payload), and by 19F-MRI at higher temperature when the release takes place (Fig. 7). Interestingly, the CEST contrast disappears when the SR is released because this contrast can be observed only as long as the paramagnetic agent is confined in the nanovesicle, condition that, on the contrary, prevents the

Fig. 7. Left: CEST effect (open squares) and 19F NMR signal (filled circles) intensity of TSLs containing a paramagnetic SR and a 19F tracer. Right: CEST- and 19F-MR images of TSLs (clinical 3.0 T scanner). The CEST contrast vanished when temperature approached the transition temperature of the liposome bilayer, while 19F signal appeared. Adapted from [25].

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detection of the 19F signal that is extensively broadened by the presence of the highly concentrated paramagnetic agent. Therefore, the release of the encapsulated payload simultaneously yields loss of CEST contrast and generation of 19F contrast. A further example relevant to the set-up of theranostic applications with liposomes loaded with paramagnetic agents has been reported by Delli Castelli et al., who investigated the MRI multicontrast properties (T1, T2, and CEST) of non targeted liposomes encapsulating the clinically approved agent Gd-HPDO3A (producing T1 and T2 contrasts) or its Thulium(III) analog Tm-HPDO3A (acting as SR and able to generate both T2 and CEST contrasts) [26]. The relevance of monitoring (almost simultaneously) different types of MRI contrast relies on the fact that each acquisition mode is differently influenced by the variables that characterize the biological microenvironment hosting the liposomes, like the tissue compartmentalization of the vesicles and/or the water dynamics across tissue membranes. This option greatly extends the informative content of the experiment, thereby offering a powerful imaging tool to shed light on the in vivo intratumor trafficking of liposomes. The time evolution of the three contrast modalities after intratumor injection of the paramagnetic liposomes (Fig. 8) was properly analyzed according to a multi-step consecutive kinetic model. The results obtained were consistent with the hypothesis that the liposomes injected in the tumor were first internalized by cells (mostly tumor associated macrophages) through a passive endocytic pathway. In a second step, the liposomes released their content (i.e. the paramagnetic agent) in the cytosol and then back to the extracellular space by exocytosis. Finally, the agent is rapidly washed out by tumor vasculature. The reliability of the kinetic model was confirmed by injecting paramagnetic pH sensitive liposomes that are known to be more prone to release their content in the endosomal (acidic) environment. Interestingly, the comparison between the kinetic rates obtained for the two liposome formulations only differed for the release step

Fig. 8. Left: top: parametric MRI maps showing the time dependence of the three contrast types (T1, T2, and CEST) assessed after intratumor injection of paramagnetic stealth liposomes (tumors are labeled by circles). T1 response was obtained from Gd-loaded liposomes, while T2 and CEST effects refer here to mice injected with Tm-loaded liposomes. Bottom left: temporal evolution of the three contrast modality (R2 = 1/T2). The cell internalization of the tumor-injected liposomes can be inferred by the decrease of T1 and CEST signals (which are more sensitive to compartmentalization effects). The release of the liposome content to the cytosol produced a late enhancement of the T1 contrast. The decrease of T1 and R2 after 24 h post-injection is consistent with the washout of the agent from the tumor. Remind that CEST effect from liposomes can be detected as long as the nanovesicles remain intact (a CEST effect of 10% can be considered within the error for the acquisition conditions used). Bottom right: schematic representation of the kinetic model used for the analysis of the data reported on the left. Adapted from [26].

that resulted to be almost one order of magnitude faster for pH sensitive nanovesicles. Though less clinically relevant than MRI, also fluorescence imaging may be suitable for reporting about drug release. In analogy to what was previously discussed for paramagnetic agent entrapped in liposome core, also the encapsulation of a fluorescence dye may lead to a quenching of the imaging response, which is here uniquely dependent on the concentration of the entrapped fluorophore and not related to the water permeability of the nanovesicle. Hence, the dilution effect resulting from the release of the dye (and the associated drug) causes a fluorescence enhancement. The well-known phenomenon of quenching/de-quenching of the fluorescence is widely used for investigating the release properties of liposomes in vitro (e.g. calcein test for assessing liposomes stability), but the preclinical/clinical translation has not been extensively explored yet. A conceptually different and elegant approach has been just reported, in which the fluorescence of a multifunctional theranostic targeted nanomedicine is activated to report about the intracellular drug release [27]. The nanoconstruct was composed by a superparamagnetic iron oxide nanoparticle core (IONP) decorated with satellite CdS:Mn/ZnS quantum dots, being, therefore, both optically and magnetically detectable. In addition, quantum dots were functionalized with STAT3 inhibitor (an anticancer agent), folate residues (as targeting motif), and methoxy-polyethylene glycol as stabilizing agent (Fig. 9). The probe is optical silent due to a quenching mediated by electron/energy transfer involving IONP, folate and STAT3 agents. However, upon incubation with tumor cells overexpressing folate receptor (human breast cancer MDA-MB-231), the probe is exposed to the cytosolic reducing agent glutathione (GSH) that disassembles the nanoparticle outer surfaces, removes the fluorescence quenching and allows the visualization of the drug release. The potential of US imaging has not yet been fully exploited in the field of theranosis in spite of its wide use in many diagnostic fields. An example has been recently reported by Wang et al., who designed echogenic microdroplets (containing liquid perfluoropentane) loaded

Fig. 9. Top: schematic representation of the multifunctional targeted nanomedicine able to act as optical reporter of intracellular drug release through a GSH-mediated activation. Bottom left: an agar phantom for in vitro characterization of the theranostic agent. The phantom consisted of four stratified agar layers. From bottom to top the layers contained: MDA-MB-231 cells (human breast cancer cells overexpressing folate receptors), the nanoagent, the nanoagent internalized into the cells, and simply agar. Left: digital photograph of the phantom; right: image obtained after UV excitation of the agent. Bottom right: phase contrast (A and C) and epifluorescence (B and D) images of cells incubated with the nanoagent. Top (A and B): MDA-MB-231 cells; bottom (C and D): thymus stromal epithelial cells (TE-71, which do not express folate receptor) used as negative control. Adapted from [27].

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Fig. 10. Top: schematic representation of the aptamer-conjugated and DOX-loaded theranostic droplets. Bottom: B-mode US images measured before and after the HIFU application that triggered drug release via acoustic droplet vaporization. Adapted from [28].

with Doxorubicin and externally conjugated with aptamers (Fig. 10) [28]. The aptamers were selected to target epitopes overexpressed by a cell line of lymphoid leukemia. Application of HIFU triggered acoustic droplet vaporization (ADV), which resulted in both mechanical cancer cell destruction by inertial cavitation and chemotherapeutic treatment through the localized release of Doxorubicin. B-mode US imaging revealed contrast enhancement by ADV, thus providing imaging support to the release of the drug. 4. Monitoring therapy by imaging Differently from what was presented in the previous two sections, where drug and imaging agent need to share the same delivery carrier, the imaging contribution for predicting and monitoring therapeutic outcomes does not require such a quite stringent requisite. In fact, any imaging modality able to provide anatomic or functional diagnostic information, even without administering a contrast agent, is, in principle, suitable to provide insights about the efficacy of the administered therapy. Up to now, most of the clinical diagnostic protocols are based on a morphological evaluation of a lesion but the new developments of molecular imaging protocols able to provide a functional/cellular/molecular assessment of the pathology, are opening new horizons in this field. The development of imaging protocols aimed at establishing the efficacy of a given therapeutic treatment represents the basis of personalized medicine and patient stratification, both expected to improve considerably the healthcare in the forthcoming years. Generally, these

improved protocols need the use of specific imaging reporters, and theranostic nanomedicines represent an excellent option. Imaging of drug-delivery, by its own, can be an important tool to predict the efficacy of a therapy through the assessment of effective delivery/accumulation of the drug at the target site [29]. Another example, close to clinical translation, is represented by the development of imaging reporters able to assess the in vivo distribution of specific receptors associated with a given pathology. For instance, the human epidermal growth factor receptor 2 (HER2) is upregulated in several cancer phenotypes characterized by high aggressiveness tumors, but it is limited or absent in normal adult tissue. Hence, the evaluation of the presence of HER2 receptors in patients is very important for selecting the correct therapy, considering that in the recent years several HER2-targeted treatments have been developed, and, in fact, diagnosis of HER2 overexpression is recommended for all newly diagnosed breast carcinomas. However, approximately 10–20% of current HER2 assessments may be inaccurate. Thus, some eligible patients do not receive HER2-targeted therapy, and, conversely, others are exposed to unnecessary therapy. Furthermore, discordance in HER2 expression between the primary tumor and metastases complicates diagnosis of recurrent disease. Thus, a complementary in vivo diagnostic method might be beneficial for improved prediction and monitoring of therapy response. Feldwish and co-workers developed affibody molecules, i.e. small non-immunoglobulin-affinity proteins, as molecular tracers for profiling HER2 receptors in patients by PET or SPECT modalities [30]. The first administration in humans of 111In-ABY-002 and 68Ga-ABY-002 in

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patients with recurrent breast cancer indicated a good overall tolerability of such agents. Radiolabeled ABY-002 detected 9 out of 11 metastases positive to 18F-FDG as early as 2–3 h after injection, thus allowing the evaluation of the HER2 status of the lesions, which is a fundamental information for the successive patient stratification (Fig. 11). Imaging can also provide an important aid for monitoring cellular and regenerative therapies. Much attention has been devoted to the in vivo tracking of stem cells from the injection site to the target tissue/ organ due to the paramount importance of visualizing the cell homing to predict the success of the treatment. MRI is certainly the imaging modality of choice for such a purpose and superparamagnetic iron oxide nanoparticles have been shown to be very efficient for labeling cells and monitoring their fate in vivo [31]. Figdor and co-workers showed that cells labeled with such magnetic nanoparticles can be monitored and tracked in human patients [32]. Dendritic cells were incubated with SPIO nanoparticles (200 μgFe/mL) to make them detectable by MRI. Next, the cells were also labeled with 111 In-oxine for their tracking with the more sensitive (but less spatially resolved) SPECT imaging modality. Patients suffering of stage-III melanoma were subjected to imaging before and after injection of 7.5 million of labeled cells into a specific lymph node (inguinal, axillary or cervical, respectively). MRI allowed a very precise localization of the injection site (Fig. 12 top), thereby providing clinicians with a powerful imaging tool to rapidly assess the spatial accuracy of the cell delivery (Fig. 12 middle). MRI also offered an excellent imaging guidance to monitor cell migration from the injected site to neighboring lymph nodes (Fig. 12 bottom). The technique resulted quite sensitive and a detection limit of 1.5 × 10 5 cells was found in this experimental clinical setup. Importantly, when the cell delivery was successful, the output from MRI was much better than SPECT in terms of number of imagingdetected regional lymph nodes addressed by the labeled cells migrated. 5. Theranostic agents for radiation-based therapies The use of theranostic agents is not limited to chemotherapy, but they can have a relevant role to guide radiation-based therapies like photodynamic therapy (PDT), hyperthermia, and radiotherapy.

Fig. 11. PET/CT clinical study of a patient administered with the same radio-dose (ca. 270 MBq) of 18F-FDG and 68Ga-ABY-002. A potential metastasis (arrows) in chest wall near the axilla is seen with 68Ga-ABY-002 (right) on transverse PET (top), CT (middle), and fusion PET/CT (bottom) images. This metastasis was not visible with 18F-FDG. Adapted from [30].

Fig. 12. Top: in vivo MRI images of a human patient suffering of a stage-III melanoma before (left) and after (right) injection of SPIO-labeled dendritic cells. White arrows indicate the injection lymph node. The darkening observed in the right image demonstrates a successful cell delivery. Middle: monitoring the delivery accuracy. MRI images were taken before (left) and after (right) cell injection. In this specific case, cells were not delivered successfully (white arrow: injected cells, black arrow: target lymph node). Bottom: monitoring cell migration. MRI image 2 days after the delivery of labeled cells. Arrows indicate the injection lymph node (1), and two neighbor lymph nodes (2 and 5) where cells migrated. Adapted from [32].

In photodynamic therapy (PDT), a photosensitizer needs to be specifically delivered to the pathological cells that, upon light activation, generate singlet oxygen and reactive oxygen species (ROS) that induce apoptosis. In order to visualize the accumulation of the “killer” photosensitizer molecule, the functionalization with an imaging reporter is often necessary. Furthermore, to improve selectivity and therapeutic efficacy, tumor targeted nanoparticles are being explored in virtue of their ability of delivering a high number of photosensitizer moieties, thus making easier the loading of the imaging agent. A different, and very elegant approach is to design smart systems in which both the photosensitizer and the imaging probe are activated in the pathological district by a specific biomarker of the disease. According to this line, Liu et al. investigated the performance of molecular beacons (PPMMPB, Fig. 13) in which the photodynamic and imaging responses are both activated by matrix metalloproteinases (MMPs) that are overexpressed in many diseases [33]. In vitro and in vivo imaging experiments carried out on a mouse model of metastatic spinal lesions of breast cancer demonstrated the MMP specific activation of the theranostic agent and its ability to differentiate tumor from healthy tissue, thus allowing a selective induction of apoptosis. Hyperthermia is another physical treatment for combating cancer that involves the heat-induced killing of tumor cells. Hyperthermia leads to apoptotic cell death caused by heating of surrounding tissues

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Fig. 13. Top: schematic representation of the photodynamic molecular beacon (PPMMPB) activated by vertebral metastases. The beacon accumulates in tissue but remains photodynamically and optically silent until cleaved by MMPs at sites of vertebral metastases, which restores both its fluorescence for imaging and generation of singlet oxygen after PDT treatment. Bottom: in vivo fluorescence images in the subcutaneous xenograft tumor model: (A) before (i) and 17 h after (ii) i.v. injection of PPMMPB and (B) PPMMPB (left tumor) versus PPscrambledB (right tumor) images after 10 min (i), 6 h (ii), and 24 h (iii) post intratumor injection. Adapted from [33].

or cells to a temperature of 42–46 °C; above 46 °C the treatment may cause undesirable necrosis of the surrounding cells. Several metal-based (Au, Ag, Fe, Co) nanoparticles are under scrutiny as theranostic agents for hyperthermia. Gold nanoparticles are the most used systems in virtue of the peculiar interaction of gold with light, which results in surface plasmon resonance (SPR). In the presence of oscillating electromagnetic radiations at a particular resonance frequency, free electrons in gold undergo SPR oscillations that can decay via heat emission. SPR frequency can be modulated upon changing size, shape, and refractive index of the material. Ideally, this frequency is tuned in the near-infrared region in order to improve the light penetration into tissues. To combine imaging and therapy in the case of goldcontaining nanoparticles, molecular architectures with iron oxide cores and gold shells have been widely explored. The desired NIR absorption frequency for gold nanoparticles can be achieved by coating them with silica, which has a high dielectric constant and helps to lower the absorption frequency toward NIR wavelengths. Iron oxide particles, in addition to provide contrast in MR images, can be exploited for guiding the theranostic agent to the site of interest by means of an externally applied magnetic field (magnetic field-directed hyperthermia). As representative example of this approach, we report the work of Lim et al. who developed nanoparticles consisting of a 10 nm iron oxide core covered with 2–3 nm hollow gold nanoshells (HGNS) [34]. The surface of the HGNS particles was functionalized with an anti-HER2 monoclonal antibody and with a fluorophore (Fig. 14). The specificity of the antibody targeting ability was tested in SKBR3 (HER2-positive) and MCF-7

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(HER2-negative) cancer cells. Fluorescence was only detected in SKBR3 cells, thus implying that a successful HER2 uptake of the HGNS nanoparticles. Photothermal therapy performed on treated SKBR3 cells with an 808 nm laser resulted in cell death within 3 min of exposure without causing an appreciable damage to unexposed cells. Radiotherapy can be applied externally or internally, the latter occurring through a dosage of radioactive materials which emit radiation upon their decay, thus causing cell death due to irreversible DNA damage and apoptosis induced by the generation of free radicals. If employed without any cell surface targeting groups, “internal” radiation therapy does not have any specificity toward cancer cells over normal cells. An approach to specifically deliver a radiation dose relies on the administration of radioactive containing peptides (radiopeptides). Radiopeptides are attractive options in the field of cancer therapeutics because they have rapid clearance, fast tissue penetration, good target accessibility, low antigenicity, ease of synthesis, and they are a good option in case of non-operable or metastasized tumors. Besides the therapeutic β-emitting isotopes (e.g. yttrium-90 or lutetium-177), radiopeptides can be also labeled with γ- (e.g. technetium-99 m) or positronemitting (e.g. fluorine-18) nuclides that allow their detection via PET or SPECT modalities. In particular, it has been reported that radiopeptides targeting somatostatin receptors (overexpressed in the majority of neuroendocrine malignancies) have a great potential for both imaging and therapy of tumors where other therapies fail. The development of gallium-68-labeled somatostatin analogs such as DOTA-NOC (DOTA-1Nal3-octreotide), DOTA-TOC (DOTA-D-Phe1-Tyr3-octreotide), or DOTATATE (DOTA-D-Phe1-Tyr3-Thr8-octreotide) for PET/CT imaging has dramatically improved the diagnosis of neuroendocrine tumors. Somatostatin analogs containing the FDA-approved ligands DTPAoctreotide, DOTA-octreotide, DTPA-Tyr3-octreotide/octreotate, and DOTA-Tyr3-octreotide/octreotate have high binding affinity for the subtype 2 (sst2) of somatostatin receptors. Fig. 15 shows PET/CT images of a child suffering from a neuroblastoma tumor imaged with 68Ga-DOTATATE [35]. Such imaging response, indicating an high expression of somatostatin receptors, is of paramount importance for guiding radiotherapy carried out with the

Fig. 14. Top: overview of the preparation of the antibody-coated iron-oxide-containing gold nanoshells. Bottom: SKBR3 cells incubated with the targeted nanotheranostic agent and stained with calcein-AM (green). Irradiation with an 808 nm laser resulted in a significant decrease in survival of treated cells. Adapted from [34].

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Fig. 16. Top left: schematic structure of the Gd-B-L theranostic agent. Top right: correlation between the boron uptake determined by MRI versus the amount measured ex-vivo by ICP-MS for two tumor cell lines (B16, filled squares, and HepG2, open squares). Bottom left: T1w-MR images of B16 melanoma-bearing mouse before (PRE) and 4 h after administration of Gd/B/LDL particles. Bottom right: T2w-MR images acquired on a mouse irradiated with neutrons without (left) or with (right) administration of Gd/B/LDL. Images were acquired 12 days after neutron irradiation. White arrows indicate the tumor lesion. Adapted from [36]. Fig. 15. Top right (A): PET/CT images showing 68Ga-DOTATATE–avid lesions in T4 vertebral body and three liver metastases (arrow). Physiologic uptake is seen in pituitary, kidneys, bladder, stomach wall, liver, and spleen. Bottom right (B): images taken on the same patient after 3 administrations of 177Lu-DOTATATE, showing metabolic partial response with lesion reduction for T4 and liver and no new lesions. Cross-sectional images on the left show reduction in size of liver lesions. Adapted from [35]. 177

Lu-analog of the same chelator. A 68Ga PET/CT scan, taken after three administrations of the radiotherapeutic agent, highlighted a positive therapeutic outcome with a regression in some lesions and an apparent block of the metastatic activity. Another important therapeutic scenario where imaging could help planning and monitoring of radiation-based therapies is represented by BNCT (Boron Neutron Capture Therapy). BNCT is a therapeutic modality based on the selective nuclear reaction that occurs when nonradioactive 10B nuclei are irradiated with low-energy thermal neutrons to yield very toxic α particles. The advantages of this treatment relies on the high nuclear capture cross section of 10B nuclei, and on the fact that cytotoxic effects are confined to the single cell in which the BNCT agent has been entrapped. The selective uptake by tumor cells is one of most challenging issues for a successful BNCT treatment. In addition, very high doses of the boron-containing agent are required (ca. 1 billion 10B nuclei/cell). There is the necessity to use a non-invasive method that could quantify the concentration of B atoms in the tumor, thus improving safety and optimizing the therapeutic outcome. Recently, Geninatti et al. demonstrated the potential of MRI to report about the delivery of a BNCT agent in a mouse tumor model [36]. They synthetized a Gd-based agent (Fig. 16) functionalized with a carborane moiety and an aliphatic tail aimed at favoring the formation of adduct with low density lipoproteins (LDLs), known to be avidly taken up by tumor cells that upregulate LDL receptors. The cellular uptake of Gd-B-LDL construct was demonstrated first in cellulo, where a very good correlation between the amount of boron internalized and the MRI response generated by the Gd-agent was observed (Fig. 16). Then, the theranostic agent was tested in vivo on a syngeneic grafted melanoma B16 mouse model. MR images displayed in Fig. 15 highlight the detection of the delivery of the agent in the tumor. Furthermore, the tumor uptake of a large amount of boron nuclei allowed a significant regression of the tumor lesion after exposure to a proper neutron flux. 6. Imaging-guided surgery This overview concludes with a recently reported application of imaging probes in imaging-guided surgery. Most of the clinical

interventional procedures are guided by human vision and perception. However, human eyes are not sensitive or accurate enough in detecting the cellular or molecular signature of a given disease. For this reason advanced optical and opto-acoustic methods have been considering to complement human vision for making clinical decision during interventions [37]. A variety of optical-based techniques have emerged with the potential to provide intra-operatively guidance especially for cancer surgery in order to support the surgeon for a reliable and accurate evaluation of tumor rims. The great advances achieved in this field have

Fig. 17. Top: color image (A) with the corresponding tumor-specific fluorescence image (B) of a representative area in the abdominal cavity. Bottom: scoring of the number of detected tumor spots was the result obtained from the inspection of five independent surgeons. Adapted from [38].

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been made possible by the important technological advances that have allowed the transfer of the necessary optical apparatus to the surgery room. Fig. 17 reports an example in which a much more accurate intra-operative counting of tumor spots was guided by fluorescence. The optical signal was generated by a fluorescein isothiocyanate (FITC)based dye properly designed to target folate-receptor α (FR-α), which is overexpressed in 90–95% of the epithelial ovarian cancer phenotype [38]. This work has attracted much attention and it outlines the great clinical benefits of intraoperative tumor-specific optical imaging upon the systemic administration of a targeted fluorescent dye. 7. Concluding remarks Imaging-guided therapy is growing fast and appears to be one of the most important products of the tremendous achievements attained in the fields of imaging technology and molecular imaging. The role of imaging in pursuing improved therapeutic treatments is further stressed by the outstanding results recently obtained in the field of imaging-guided surgery. For the future, one may easily foresee that “theranostics” will have an important role in the development of personalized medicine protocols. More work appears necessary to develop smart systems that couple excellent targeting capabilities with optimal drug-release responsiveness either to internal or external stimuli. To tackle this task, it is even more important now than in the past to pursue an improved integration of the involved teams with contributions from chemistry, biology, medicine, and imaging/therapy engineers. Acknowledgments This research was supported by funding from the University of Torino (code D15E11001710003), Regione Piemonte (PIIMDMT and Nano-IGT projects), MIUR (PRIN 2009), EU-FP7 integrated project ENCITE and ESF COST Action TD1004 (Theranostics Imaging and Therapy: An Action to Develop Novel Nanosized Systems for ImagingGuided Drug Delivery). Scientific support from CIRCMSB (Consorzio Interuniversitario di Ricerca sulla Chimica dei Metalli nei Sistemi Biologici) is also gratefully acknowledged. References [1] J.L. Arias, B. Clares, M.E. Morales, V. Gallardo, M.A. Ruiz, Lipid-based drug delivery systems for cancer treatment, Curr. Drug Targets 12 (2011) 1151–1165. [2] P. Tanner, P. Baumann, R. Enea, O. Onaca, C. Palivan, W. Meier, Polymeric vesicles: from drug carriers to nanoreactors and artificial organelles, Acc. Chem. Res. 44 (2011) 1039–1049. [3] W.J. Stark, Nanoparticles in biological systems, Angew. Chem. Int. Ed Engl. 50 (2011) 1242–1258. [4] A.L. Laza-Knoerr, R. Gref, P. Couvreur, Cyclodextrins for drug delivery, J. Drug Target. 18 (2010) 645–656. [5] K.K. Ng, J.F. Lovell, G. Zheng, Lipoprotein-inspired nanoparticles for cancer theranostics, Acc. Chem. Res. 44 (2011) 1105–1113. [6] F.J. Kratz, Albumin as a drug carrier: design of prodrugs, drug conjugates and nanoparticles, J. Control. Release 132 (2008) 171–183. [7] J.W. Yoo, D.J. Irvine, D.E. Discher, S. Mitragotri, Bio-inspired, bioengineered and biomimetic drug delivery carriers, Nat. Rev. Drug Discov. 10 (2011) 521–535. [8] S. Ganta, H. Devalapally, A. Shahiwala, M. Amiji, A review of stimuli-responsive nanocarriers for drug and gene delivery, J. Control. Release 126 (2008) 187–202. [9] R. Weissleder, Molecular imaging: exploring the next frontier, Radiology 212 (1999) 609–614. [10] X. Chen, S.S. Gambhir, J. Cheon, Theranostic nanomedicine (editorial), Acc. Chem. Res. 44 (2011) 841. [11] C.T. Moonen, Spatio-temporal control of gene expression and cancer treatment using magnetic resonance imaging-guided focused ultrasound, Clin. Cancer Res. 13 (2007) 3482–3489. [12] Z. Li, Q. Jin, C. Huang, S. Dasa, L. Chen, L.P. Yap, S. Liu, H. Cai, R. Park, P.S. Conti, Trackable and targeted phage as positron emission tomography (PET) agent for cancer imaging, Theranostics 1 (2011) 371–380. [13] E. Terreno, D. Delli Castelli, A. Viale, S. Aime, Challenges for molecular magnetic resonance imaging, Chem. Rev. 110 (2010) 3019–3042.

337

[14] W.J. Mulder, G.J. Strijkers, G.A. van Tilborg, A.W. Griffioen, K. Nicolay, Lipid-based nanoparticles for contrast-enhanced MRI and molecular imaging, NMR Biomed. 19 (2006) 142–164. [15] C. Grange, S. Geninatti-Crich, G. Esposito, D. Alberti, L. Tei, B. Bussolati, S. Aime, G. Camussi, Combined delivery and magnetic resonance imaging of neural cell adhesion molecule-targeted doxorubicin-containing liposomes in experimentally induced Kaposi's sarcoma, Cancer Res. 70 (2010) 2180–2190. [16] E. Cittadino, M. Ferraretto, E. Torres, A. Maiocchi, B.J. Crielaard, T. Lammers, G. Storm, S. Aime, E. Terreno, MRI evaluation of the antitumor activity of paramagnetic liposomes loaded with prednisolone phosphate, Eur. J. Pharm. Sci. 45 (2012) 436–441. [17] F.M. Kievit, M. Zhang, Surface engineering of iron oxide nanoparticles for targeted cancer therapy, Acc. Chem. Res. 44 (2011) 853–862. [18] K. Kaaki, K. Hervé-Aubert, M. Chiper, A. Shkilnyy, M. Soucé, R. Benoit, A. Paillard, P. Dubois, M.-L. Saboungi, I. Chourpa, Magnetic nanocarriers of doxorubicin coated with poly(ethylene glycol) and folic acid: relation between coating structure, surface properties, colloidal stability, and cancer cell targeting, Langmuir 28 (2012) 1496–1505. [19] J.R. McCarthy, E. Korngold, R. Weissleder, F.A. Jaffer, A light-activated theranostic nanoagent for targeted macrophage ablation in inflammatory atherosclerosis, Small 6 (2010) 2041–2049. [20] D. Delli Castelli, E. Gianolio, S. Geninatti Crich, E. Terreno, S. Aime, Metal containing nanosized systems for MR-molecular imaging applications, Coord. Chem. Rev. 252 (2008) 2424–2443. [21] E. Terreno, A. Sanino, C. Carrera, D. Delli Castelli, G.B. Giovenzana, A. Lombardi, R. Mazzon, L. Milone, M. Visigalli, S. Aime, Determination of water permeability of paramagnetic liposomes of interest in MRI field, J. Inorg. Biochem. 102 (2008) 1112–1119. [22] B.L. Viglianti, S.A. Abraham, C.R. Michelich, P.S. Yarmolenko, J.R. MacFall, M.B. Bally, M.W. Dewhirst, In vivo monitoring of tissue pharmacokinetics of liposome/drug using MRI: illustration of targeted delivery, Magn. Reson. Med. 51 (2004) 1153–1162. [23] A.H. Negussiey, P.S. Yarmolenko, A. Partanen, A. Ranjan, G. Jacobs, D. Woods, H. Bryant, D. Thomasson, M.W. Dewhirst, B.J. Wood, M.R. Dreher, Formulation and characterisation of magnetic resonance imageable thermally sensitive liposomes for use with magnetic resonance-guided high intensity focused ultrasound, Int. J. Hypethermia 27 (2011) 140–155. [24] M. de Smet, E. Heijman, S. Langereis, N.M. Hijnen, H. Grüll, Magnetic resonance imaging of high intensity focused ultrasound mediated drug delivery from temperature-sensitive liposomes: an in vivo proof-of-concept study, J. Control. Release 150 (2011) 102–110. [25] S. Langereis, J. Keupp, J.L.J. van Velthoven, I.H.C. de Roos, D. Burdinski, J.A. Pikkemaat, H. Gruell, A temperature-sensitive liposomal 1H CEST and 19F contrast agent for MR image-guided drug delivery, J. Am. Chem. Soc. 131 (2009) 1380–1381. [26] D. Delli Castelli, W. Dastrù, E. Terreno, E. Cittadino, F. Mainini, E. Torres, M. Spadaro, S. Aime, In vivo MRI multicontrast kinetic analysis of the uptake and intracellular trafficking of paramagnetically labeled liposomes, J. Control. Release 144 (2010) 271–279. [27] R.N. Mitra, M. Doshi, X. Zhang, J.C. Tyus, N. Bengtsson, S. Fletcher, B.D.G. Page, J. Turkson, A.J. Gesquiere, P.T. Gunning, G.A. Walter, S. Santra, An activatable multimodal/multifunctional nanoprobe for direct imaging of intracellular drug delivery, Biomaterials 33 (2012) 1500–1508. [28] C.-H. Wang, S.-T. Kang, Y.-H. Lee, Y.-L. Luo, Y.-F. Huang, C.-K. Yeh, Aptamer-conjugated and drug-loaded acoustic droplets for ultrasound theranosis, Biomaterials 33 (2012) 1939–1947. [29] T. Lammers, S. Aime, W.E. Hennink, G. Storm, F. Kiessling, Theranostic nanomedicines, Acc. Chem. Res. 44 (2011) 1029–1038. [30] R.P. Baum, V. Prasad, D. Muller, C. Schuchardt, A. Orlova, A. Wennborg, V. Tolmachev, J. Feldwisch, Molecular imaging of HER2-expressing malignant tumors in breast cancer patients using synthetic 111In- or 68Ga-labeled affibody molecules, J. Nucl. Med. 51 (2010) 892–897. [31] M. Modo, M. Hoehn, J.W. Bulte, Cellular MR imaging, Mol. Imaging 4 (2005) 143–164. [32] J.M. de Vries, W.J. Lesterhuis, J.O. Barentsz, P. Verdijk, J.H. van Krieken, O.C. Boerman, W.J.G. Oyen, J.J. Bonenkamp, J.B. Boezeman, G.J. Adema, J.W.M. Bulte, T.W.J. Scheenen, C.J.A. Punt, A. Heerschap, C.G. Figdor, Magnetic resonance tracking of dendritic cells in melanoma patients for monitoring of cellular therapy, Nat. Biotechnol. 23 (2005) 1407–1413. [33] T.W. Liu, M.K. Akens, J. Chen, L. Wise-Milestone, B.C. Wilson, G. Zheng, Imaging of specific activation of photodynamic molecular beacons in breast cancer vertebral metastases, Bioconjug. Chem. 22 (2011) 1021–1030. [34] Y.T. Lim, M.Y. Cho, J.K. Kim, S. Hwangbo, B.H. Chung, Plasmonic magnetic nanostructure for bimodal imaging and photonic-based therapy of cancer cells, Chembiochem 8 (2007) 2204–2209. [35] J.E. Gains, J.B. Bomanji, N.L. Fersht, T. Sullivan, D. D'Souza, K.P. Sullivan, M. Aldridge, W. Waddington, M.N. Gaze, 177Lu-DOTATATE molecular radiotherapy for childhood neuroblastoma, J. Nucl. Med. 52 (2011) 1041–1047. [36] S. Geninatti-Crich, D. Alberti, I. Szabo, A. Deagostino, A. Toppino, A. Barge, F. Ballarini, S. Bortolussi, P. Bruschi, N. Protti, S. Stella, S. Altieri, P. Venturello, S. Aime, MRI-guided neutron capture therapy by use of a dual gadolinium/boron agent targeted at tumour cells through upregulated low-density lipoprotein transporters, Chemistry 17 (2011) 8479–8486. [37] A. Sarantopoulos, N. Beziere, V. Ntziachristos, Optical and opto-acoustic interventional imaging, Ann. Biomed. Eng. (2012), http://dx.doi.org/10.1007/s10439-011-0501-4. [38] G.M. van Dam, G. Themelis, L.M. Crane, N.J. Harlaar, R.G. Pleijhuis, W. Kelder, A. Sarantopoulos, J.S. de Jong, H.J. Arts, A.G. van der Zee, J. Bart, P.S. Low, V. Ntziachristos, Intraoperative tumor-specific fluorescence imaging in ovarian cancer by folate receptor-α targeting: first in-human results, Nat. Med. 17 (2011) 1315–1319.