*1.2. Nanoparticles*

In its first applications, hybrid PET/MRI was realized through the simultaneous administration of a mixture of MRI and PET probes, resulting in a cocktail of imaging agents causing high risk for the patient [38]. Additionally, this mixture could not guarantee an exact spatial and temporal correlation of the two imaging modalities due to the different biodistribution, pharmacodynamic and pharmacokinetic properties of the imaging agent. To overcome these limitations, nanoparticles (NPs) were proposed as delivery systems for different imaging agents to obtain bimodal probes for the simultaneous monitoring of both modalities. NPs are defined as particles with at least one dimension lying between 1 and 100 nm [39,40]. In recent years, very different NPs such as proteins, polymers, dendrimers, micelles, liposomes, viral capsids, metal oxides (iron-oxide NPs), zeolites, and mesoporous silicas were investigated, and very different shapes, such as spheres, cylinders (nanorods), and tubes were explored [24].

An NP-based PET/MRI bimodal probe constitutes three essential components: a carrier, a PET tracer usually represented by a positron emitter radioisotope characterized by high sensitivity (e.g., 18F-fluorodeoxyglucose), and an MRI component (e.g., gadopentetic acid (Gd-DTPA)), providing high tissue contrast and resolution. The MRI component can either work as a carrier itself (iron-oxide or gadolinium-oxide NPs) or can be a moiety bound to or entrapped in the carrier (e.g., Gd ions grafted onto NPs and polymeric matrices or biologically derived nanosized systems like apoferritin cages [41] and low-density lipoprotein (LDL) particles [42], respectively). Some possible configurations are reported in Figure 1.

As carriers, NPs offer a number of different design options, and the tailoring of their properties can be exploited to directly impact the in vivo fate of the resulting probe. Particle size, charge, core and surface properties, shape, and multivalency are the main features to be finely tuned in order to achieve a proper in vivo distribution, confer a targeting ability, and reduce toxicity of the NPs [43]. The hydrodynamic size determines the NP fate in the body, since vectors with a mean diameter smaller than 5 nm are usually eliminated by renal excretion, whereas larger particles (100 nm) are easily taken up by macrophages [44,45]. NP shape influences the internalization into cells that is relevant in cell tracking and labeling; for example, rod-like particles present higher internalization rates compared to spherical particles [46]. This phenomenon can be explained considering its similarity to rod-like bacterium internalization in nonphagocytic cells [47].

**Figure 1.** Multimodal nanoparticles. (**A**) Multimodal nanoparticle composed by a core (representing the magnetic resonance imaging (MRI) component) and a shell functionalized with an antibody. The positron emission tomography (PET) radiotracer is chelated and bound to the spacer. (**B**) A polymeric nanoparticle entrapping paramagnetic moieties is represented, where the PET radiotracer is chelated and bound to the spacer. (**C**) Liposomal formulation entraps paramagnetic moieties in the aqueous inner core, while the PET component is covalently linked to the spacer. (**D**) Liposomal formulation with paramagnetic ion inserted in the bilayer.

After NP injection into the bloodstream, they are rapidly coated by plasma proteins in a process called opsonization. The NPs are then recognized by plasma membrane receptors found on monocytes and macrophages and are, thus, taken up by the body's main defense system, the reticuloendothelial system (RES), also known as the mononuclear phagocyte system (MPS). The liver, spleen, and bone marrow are rich in macrophages, thus becoming the most accessible organs to NPs [48]. For these reasons, NPs should be coated by adequate materials, to avoid nonspecific uptake by the RES [48–50] (stealth effect). In general, hydrophilic and neutral surfaces do not tend to interact with blood components (serum proteins); therefore, they are optimal for minimizing opsonization and clearance [51,52]. Since neutral polymers have no functional groups (amine, carboxyl, or hydroxyl) for ligand linkage, a further step of functional group activation is often mandatory. Another coating strategy employs hydrophilic bifunctional materials such as biphosphonate [53] or aluminum hydroxide [54]. NPs with biocompatible coating layers such as polymers (polyethylene glycol (PEG)), dendrimers, polysaccharides (dextran and chitosan), and polypeptides (serum albumin) can have enhanced properties including better stability in terms of agglomeration, biocompatibility, and solubility in water, along with low toxicity. The most used coating polymers are dextran, chitosan and, above all, PEG [41]. PEG is a hydrophilic, water-soluble, biocompatible polymer widely used to reduce opsonization and increase circulation time from seconds or minutes up to hours [55]. It is important to notice that surface modifications may have an impact on the superparamagnetic properties of iron-oxide NPs; for this reason, coating materials must be carefully chosen [56]. In particular, the nature and the thickness of the coating affect relaxivity. A more hydrophilic coating material results in more water molecules being retained for interacting with the magnetic centers; on the other hand, a thicker coating results in more protons being shielded from the magnetic field [57].

NP delivery to malignant cells can be achieved through both passive and active targeting. Passive targeting is due to the enhanced permeability and retention effect (EPR); since tumor vessels have larger fenestrations, the vascular permeability is higher, and NPs can easily extravasate in tumor tissue. Moreover, the inefficient lymphatic drainage contributes to NP retention in the tumor interstitial space [58]. Even though non-targeted NPs can accumulate in the tumor region due to the EPR effect, the lack of efficient lymphatic drainage generates an increase in interstitial pressure and, consequently, a drop in pressure gradient between the vessel and the extracellular space, causing nanoparticle stacking around the vessel wall [44,59]. For these reasons, there is a need for the development of NPs capable of efficiently and specifically targeting tumor cells [27]. The high NP surface-to-volume ratio helps to overcome this limitation since the NP surface can be functionalized through target-specific moieties that allow an active targeting of cancer cells.

Finally, multivalence refers to the ability to bind different imaging probes, targeting ligands, and therapeutic formulations. This feature is very important for multimodal and molecular imaging where a significant number of targeting probes are needed to track a specific biological path.

#### *1.3. Radiolabeled Nanoparticles*

Tracers currently used in clinical practice are labeled using positron emitters with a relatively low half-life time ranging from 2.037 to 109.8 min. Most of the radionuclides used for labeling are produced via a cyclotron. It generates a beam of accelerated protons and deuterons that are used to irradiate a target (e.g., 14N2 gas, 20Ne gas, 18O water or gas), thereby giving the desired radioisotope through a nuclear reaction. Table 3 shows the main radionuclides with related half-life time, the average energy of positron (β+), and means of production.


**Table 3.** Principal radionuclides and related features.

\* The values were obtained from the database of the National Nuclear Data Center (NNDC) at Brookhaven National Laboratory, Upton NY, USA. \*\* Mean energy of the β spectrum.

In PET clinical applications, 18F is one of the most suitable radionuclides for radiotracer synthesis since 97% of isotope decay is via positron emission [60], with a fairly low energy of positron emission (maximum 0.635 MeV) and an optimal half-life of 109.8 min, which is considered acceptable for chemical syntheses and favorable when investigating biological processes with a time frame longer than 100 min [61]. 18F-based radiotracers are essentially synthesized through two reactions: nucleophilic substitution or electrophilic substitution. Frequently, 18F is introduced to replace hydrogen in biomolecules. However, in terms of size, the van der Waals radius of 18F (1.47 Å) is closer to oxygen (1.52 Å) than that of hydrogen (1.20 Å) [62]; thus, 18F is generally obtained starting from water enriched with 18O through a nuclear reaction like 18O(p, n)18F.

PET radiotracers for cancer diagnosis can be grouped based on their target mechanism as follows:


Other widely used radionuclides are 68Ga and 64Cu. In particular, 64Cu is gaining increasing interest for its theranostic potential [63]; during its decay, it emits both positron and Auger electrons allowing for both PET imaging and internal targeted radiation therapy. Indeed, Auger-emitting radionuclides that localize in the nucleus of tumor cells demonstrate a potential for cancer therapy. However, their biological effect is critically dependent on their sub-cellular (and sub-nuclear) localization [64] and on the DNA topology [65].

The chemical structures of the most common radiotracers are reported in Figure 2.

NP radiolabeling with the abovementioned tracers can be achieved through different techniques. In the literature, the four following main strategies are reported [66]:


**Figure 2.** Chemical structure of fluorine-based radiopharmaceuticals.

The coordination chemistry approach is the most used since radioisotopes can be chelated by different molecules that are covalently bound directly to the NP surface. A strong linkage between the chelator coordinating the radioisotope and the NP surface is desired to assure the stability of the radiolabeling. It is worth noting that many exogenous chelators can currently only coordinate with certain radioisotopes, meaning that an effective chelator-based radiolabeling requires the selection of the best chelator for the isotope of interest [67]. In addition, the choice of the chelating agent should be such to minimize in vivo transchelation.

New chelators for metallic radioisotopes were recently synthesized, including tetradentate acyclic chelators such as PTMS, esadentate acyclic chelators such as ethylenediaminetetraacetic acid EDTA or DTPA, and macrocyclic chelators such as 1,4,7-triazacyclononane-N,N',N"-triacetic acid (NOTA) and 1,4,7,10-Tetraazacyclododecane-1,4,7,10-tetraacetic acid DOTA [68], whose chemical structures are presented in Figure 3.

Recently, Laverman et al. reported the possibility of 18F chelation through an "Al–18F" complex, which carries out a coordination bond with the macrocyclic chelator NOTA [69].

Direct bombardment is achieved by direct irradiation of inorganic NPs with protons and neutrons to obtain radiolabeled NPs. Perez-Campana et al. demonstrated the nuclear reaction 16O(p, α) 13N on Al2O3 NPs, where the radioisotope is incorporated in the inorganic NPs without any modification of the particle surface and morphology. Moreover, they demonstrated the stability of the radiolabeling by monitoring the in vivo signal after NP intravenous (i.v.) injection [70]. The main limitation of this approach is related to its application to functionalized NPs; the irradiation procedure may induce damages to the organic molecules conjugated onto the NP surface, causing the loss of their biological activity.

**Figure 3.** DOTA and NOTA chelators: chemical and three-dimensional structures. 1,4,7 triazacyclononane-N,N',N"-triacetic acid (NOTA) and 1,4,7,10-Tetraazacyclododecane-1,4,7,10 tetraacetic acid (DOTA).

An alternative approach is the synthesis of radioactive NPs starting from radioactive and non-radioactive precursors. 64Cu is the most widely used radioisotope for this strategy thanks to which both organic and inorganic NPs can be obtained as liposomal 64Cu, [64Cu] CuS, or [64Cu] CuFe3O4 [71–73]. However, high temperatures and elevated incubation times are required for their production; thus, radiocontamination problems may arise.

Finally, post-synthesis NP radiolabeling seems to be a very promising chelator-free approach. However, both NP properties and chemical and physical interactions between NPs and the radioisotope have to be carefully taken into account. As an example, Chakravarty et al. produced a probe for dual MRI/PET imaging by 69Ge radiolabeling of superparamagnetic iron-oxide NPs (SPIONs). They were realized by exploiting the unique interaction between the NP surface and the radiotracer contact, overcoming all the limitations associated with the complex 69Ge coordination chemistry of traditional chelator-based methods [74].

Moreover, by exploiting the ability of some radiotracers to emit α and β particles, radiolabeled NPs can be used for radiation therapy in theranostic applications. These radiotracers, indeed, generate ionization in the atoms (mostly in water molecules), with the formation of free radicals and consequent damage to cellular DNA. As an example, liposomes containing α-emitters are widely described in the literature for their ability to improve the radionuclide circulation time and mediate its interaction with the biological environment [75–77]. Through this approach, it is possible to improve the ratio between radiation dose to tumor and normal tissues. Secondly, because of a better time to circulate, these formulations cause larger concentrations to diffuse within the tumor tissue and may, therefore, provide a less heterogeneous tumor dose [75].

#### *1.4. PET*/*MRI Nanoparticles and Preclinical Applications*

NPs were extensively studied at a preclinical level as imaging probes for dual MRI/PET tumor imaging. According to the chemical composition of the core, NPs can be classified into inorganic and organic [78]. Inorganic NPs recently gained significant attention due to their unique physical and chemical properties. In particular, their chemical inertness, good stability, and the easiness of surface functionalization make inorganic NPs attractive for imaging of malignant tumors. However, their toxicity remains the main concern; it was demonstrated that iron-oxide NPs entering into cells through endocytosis show high toxicity because of their accumulation in endo-lysosomal compartments [59]. The most used carriers of this category are iron-oxide NPs and silica NPs. Common nanoconstructs are shown in Figure 4.

**Figure 4.** Typical nanocarriers: (from left) superparamagnetic iron-oxide nanoparticles, silica-based nanoparticles, liposomes, micelles, polymeric nanoparticles, and dendrimers.

#### 1.4.1. Iron-Oxide Nanoparticles

Magnetic iron-oxide NPs, typically magnetite Fe3O4 and maghemite, γ Fe2O3, are broadly employed in MRI imaging especially for the liver, spleen, and bone marrow due, to their ability to shorten T2 and T2\* relaxation times. According to their size, they can be categorized into *micrometer-sized paramagnetic iron oxide* (MPIO) (several micrometers), *superparamagnetic iron oxide* (SPIO) (hundreds of nanometers), *ultrasmall superparamagnetic iron oxide* (USPIO) (below 50 nm) [79], and *mono crystalline iron oxide* (MION) (representing a subset of USPIO ranging from 10 to 30 nm) [78,80].

The most common method for SPIO and USPIO synthesis is the reduction and coprecipitation reaction of ferrous and ferric salts in a basic aqueous media [81–83]. Resulting NPs are generally polydisperse and poorly crystalline; therefore, other preparation methods are often preferred, such as thermal decomposition and microwave synthesis [17]. Bare NPs are prone to agglomeration due to their high surface energy. In order to improve both colloidal and chemical stability, many polymeric coating materials were proposed, such as dextran, carboxymethylated dextran, carboxydextran, chitosan, starch, PEG, heparin, albumin, arabinogalactan, glycosaminoglycan, sulfonated styrene–divinylbenzene, organic siloxane, polyvinyl alcohol, poloxamers, and polyoxamines [84,85]. In addition, the polymeric corona is able to protect iron-oxide NPs, preventing erosion at acidic pH, lowering cytotoxicity [63]. The coating can be performed during the co-precipitation process, with the synthesis of the NPs occurring simultaneously to its coating [86,87] or post-synthesis, with the coating realized after the synthesis of the NPs [88,89]. Surface coating is a key factor for NP bioconjugation to biological ligands such as peptides or antibodies; therefore, it represents clinical potential for cancer imaging. Nevertheless, iron-oxide NPs have some important drawbacks. First of all, they act as negative contrast and, after administration, there is a loss of signal that makes medical evaluation less easy compared to T1 CA brightness. Moreover, the high susceptibility causes distortion artefacts and reduces the contrast-to-noise ratio [79]. Gd-based T1 agents are the most extensively and clinically used. Alloy materials were investigated to obtain more efficient T2 CAs because they are endowed with higher magnetic anisotropy [69], crystallinity, and relaxivity; thus, various bimetallic ferrite NPs named *magnetic engineered iron-oxide NPs*, such as CoFe2O4, MnFe2O4, and NiFe2O4, were tested [90].

There are several commercially available superparamagnetic iron-oxide NP formulations such as Feridex (Berlex" Hanover, NJ, USA), Endorem (Guerbet, Villepinte, EU), and Resovist (Schering, EU, Japan). They are mostly used for liver and spleen tumors diagnosis [91], and the coating polymers are dextran for Feridex and Endorem, and an alkali-treated low-molecular-weight carboxydextran for Resovist [92]. Many preclinical studies were conducted to assess the iron-oxide NP potential as PET/MRI probes for cancer imaging exploiting both passive targeting (for lymph node mapping) and active targeting strategies (mainly through RGD (Arg–Gly–Asp) conjugation).

Thorek and coworkers [93] prepared 89Zr radiolabeled iron-oxide NPs (ferumoxytol) to visualize the axillary and brachial lymph node drainage in healthy wild-type mice. In detail, the iron-oxide core was surrounded by a semisynthetic polysaccharide coating of polyglucose sorbitol carboxymethylether, and desferrioxamine was used as a chelator. In the same study [93], after intraprostatic administration in Hi-Myc transgenic mice bearing invasive prostatic adenocarcinoma, PET/MRI imaging delineated draining nodes in the abdomen and the inguinal region, in addition to prostatic ones.

In 2019, Madru et al. [94] proposed a new, time-efficient, chelator-free conjugation of 64Cu on PEGylated SPIONs for PET/MRI detection and localization of sentinel lymph nodes (SLNs) in C57BL/6J mice. The stability of radiolabeling up to 24 h and NP accumulation in the SLN were demonstrated through a biodistribution study. Lymph nodes metastases are important markers for cancer staging and treatment, and their localization can be useful in presurgical planning.

Xie and colleagues [95] encapsulated iron-oxide NPs, after modification with dopamine, into human serum albumin (HSA) matrices and labeled them with Cy5.5 dye and 64Cu-DOTA. NPs were injected into a U87MG xenograft mouse model; PET and NIRF imaging showed a higher signal-to-noise ratio compared to MRI because of their higher sensitivity. On the other hand, MRI scans post NP injection showed a clear inhomogeneous distribution thanks to their high spatial resolution. These findings were confirmed by histological studies. The HSA shell conferred prolonged circulation time and lower macrophage uptake rate. Such NPs are suitable for theranostic applications if co-loaded with drug molecules.

An active targeting probe was developed by Lee and coworkers [96] who conjugated RGD to 64Cu radiolabeled iron-oxide NPs. As a coating material, polyaspartic acid was chosen since it exposes both carboxyl groups interacting with NPs and amine groups useful for DOTA and RGD conjugation. Imaging was performed on a U87MG mouse model, and both PET and MRI confirmed that the accumulation of NPs was mediated by αvβ<sup>3</sup> integrin binding. Kim and colleagues [97] injected 68Ga labeled iron-oxide NPs into BALB/c nude mice bearing colon cancer (HT-29) cells, using oleanolic acid as a tumor-targeting molecule. This ligand was shown to inhibit colon cancer cell proliferation, as well as induce apoptosis and cancer cell death. Binding assays and histological studies confirmed the tumor uptake of NPs thanks to oleanolic acid affinity for HT-29 cancer cells. PET/MRI scans provided high-quality images and precise quantification of the tumor area.
