Fe₃O₄ MNS are extensively used in biomedicine because of their biocompatibility and ease of synthesis [47, 48]. The magnetic moment of superparamagnetic Fe₃O₄ MNS is dependent on the size with smaller particles producing lower magnetic moments [49]. Hence, the size of Fe₃O₄ MNS can be tuned by changing the reaction conditions such as reflux temperature, reflux time, and heating rate. The magnetic moment of Fe₃O₄ MNS was tuned from 25 to 43, 80, and 102 emu/(g Fe) with change in the size from 4, 6, 9, and 12 nm, respectively (resulting in r ₂ values of 78, 106, 130, and 218 mM⁻¹ s⁻¹) [27]. Further, the magnetic moment of Fe₃O₄ MNS can be modified by doping transition divalent metal ions (Co²⁺, Ni²⁺, Mn²⁺, and Zn²⁺). By adding different metal precursors during Fe₃O₄ synthesis, monodisperse MFe₂O₄ MNS were synthesized [38]. The metal ferrite nanostructures have an inverse spinel crystal structure composed of face-centered cubic packed lattice of oxygen atoms with octahedral sites (O_h) occupied by Fe³⁺ and M²⁺ ions and tetrahedral sites (T_d) occupied by Fe³⁺ ions (Fig. 4a). The magnetic spins of the ions at the O_h and T_h sites align opposite to each other. Hence the spins of Fe³⁺ ions at O_h and T_h cancel each other and net magnetization of MFe₂O₄ MNS is decided by magnetic moment of M²⁺ ions (Fig. 4b). The magnetization of NiFe₂O₄, CoFe₂O₄, Fe₃O₄, and MnFe₂O₄ were found to be 85, 99, 101, 110 (emu/g metal ions), respectively, depending upon the magnetic moments of the M²⁺ ions. This resulted in r ₂ values of 152, 172, 218, and 358 mM⁻¹ s⁻¹ (Fig. 4c) [38].

Interestingly, doping of nonmagnetic Zn in Fe₃O₄ and MnFe₂O₄ MNS resulted in (Zn _x Fe_1−x )Fe₂O₄ and (Zn _x Mn_1−x )Fe₂O₄ MNS, respectively, that exhibit extremely high net magnetic moment and r ₂ relaxivity [42]. The magnetization of (Zn _x Mn_1−x )Fe₂O₄ was dependent on the Zn doping level and were found to be 125, 140, 154, 166, 175, and 137 emu/g metal ions for x = 0, 0.1, 0.2, 0.3, 0.4, and 0.8, respectively, resulting in the r ₂ values of 422, 516, 637, 754, 860, and 388 mM⁻¹ s⁻¹, respectively (Fig. 5). The r ₂ value for Zn_0.4Mn_0.6Fe₂O₄ MNS is the maximum r ₂ value of reported to date for MNS [42], eight times higher than r ₂ of Feridex [36]. (Zn _x Fe_1−x )Fe₂O₄ MNS exhibited a similar trend, but the magnetization and r ₂ values were slightly lower than (Zn _x Mn_1−x )Fe₂O₄ MNS (Fig. 5) [42].

MNS based on transition metals of Fe, Co, Ni, and their alloys have higher magnetic moments than their oxide counterparts [44, 50–53]. Saturation magnetization of bulk FeCo (240 emu/g) and Fe (218 emu/g) is particularly high compared to bulk Fe₃O₄ (90 emu/g). Using the same mass of metallic MNS would then produce a far greater impact than their oxide counterparts, improving the T ₂ contrast enhancement and therapeutic efficacy of drug delivery. However, the metallic MNS carries their own set of disadvantages like chemical instability, leaching of the noniron elements, and toxicity that renders them questionable for in vivo applications [9, 53, 54]. In addition, these pure metal nanoparticles are also ferromagnetic at room temperature, rather than superparamagnetic. This means that once they are magnetized, they will remain that way regardless of whether an external magnetic field is withdrawn, resulting in an aggregation. A number of reports have been published demonstrating coatings that prevent aggregation and ensure chemical stability of metallic MNS. Options under consideration include inert metals, such as Au and Ag, peptide capping ligands, ferrites, graphite, and silica [43, 44, 52, 55–57]. For example, after graphitic shell coating, FeCo MNS were stable up to 1 month and showed very high magnetic moment (215 emu/g) and r ₂ relaxivity (644 mM⁻¹ s⁻¹), which is far superior to conventional ferrite MNS [43, 52]. Crystalline Fe₃O₄ shell was also used to protect metallic Fe MNS that resulted in M _s and r ₂ of 164 emu/g and 220 mM⁻¹ s⁻¹, respectively [44, 56]. Co MNS were coated with Au shell to provide an inert, biocompatible and stable shell with a well-known surface chemistry. Though the bulk saturation magnetization of Co is ~160 emu/g, the measured value was found ~100 emu/g after Au coating, which was still higher than Fe₃O₄ MNS (75–80 emu/g) [55].

Hybrid MNS with two or more different functional units, such as Au–Fe₃O₄, FePt–CdS, and Fe₂O₃–carbon nanotube can be synthesized through seed mediated growth. In such a heterogeneous nanostructure, each unit exhibits its unique magnetic, optical, or electronic properties [10, 58–60]. Au–Fe₃O₄ nanostructures were prepared that preserved the optical property of Au (plasmonic absorption at ~530 nm) as well as the magnetic property of Fe₃O₄ MNS (M _s  = 80 emu/g) [61]. This approach was extended to prepare semiconductor–metal alloy [62], semiconductor–metal oxide [63], and carbon nanotube–metal oxide complex [64]. MNS have been coupled with a wide range of fluorophores for multimodal imaging applications [65–67]. Lastly, T ₁/T ₂ MRI agent were prepared by conjugating Gd(III) based chelating agent with MNS [68].

PEG is a neutral and amphiphilic polymer that has been used clinically as excipients in FDA approved pharmaceutical formulations [77]. PEG coating of MNS improves their dispersion in biological media and increases blood circulation time since they are not readily recognized by the reticuloendothelial system (RES) [78]. Lutz et al. demonstrated in situ coating of PEG onto Fe₃O₄ MNS under aqueous conditions [79] while PEG grafting was achieved by single-point chemical anchoring through different functional groups including silanes [70], phosphate derivatives [80] and dopamine [81]. Peng and Sun reported ligand exchange with bifunctional PEG with dopamine [56]. Most recently, nitrodopamine has been proposed as an ultrastable chemical anchor for MNS [82]. We have developed MNS with high buffer stability by coating Fe₃O₄ with bifunctional PEG conjugated with nitrodopamine and carboxylate terminal groups [83]. The nirodopamine was covalently attached to the Fe₃O₄ surface by one end and the carboxylate group at the other end was kept open to functionalize with targeting ligands or therapeutic agents.

Dextran is a branched polysaccharide comprised of glucose subunits and is widely used for MNS coatings because of its biocompatibility and polar interactions (chelation and hydrogen bonding). Addition of dextran during synthesis of Fe₃O₄ via the co-precipitation method resulted in dextran coated Fe₃O₄ MNS [84]. Subsequent iterations of this method produced clinically approved ferumoxtran-10 (AMI-277) and ferumoxides (AMI-25) [85–90]. These two structures have cores of ~5 nm, but differ significantly in dextran coating thickness (20–40 nm vs. 80–150 nm) which results in varying blood circulation times (24 h for ferumoxtran and 2 h for ferumoxides) [35]. Since the dextran molecules are adhered nonspecifically through hydroxyl interactions with the iron oxide core, there is always a possibility of desorption [91]. In order to prevent this, the dextran polymers were chemically cross-linked on MNS surface [92]. Using this strategy, clinically approved ferumoxytol [93] and ferucarbotran [22] have been synthesized.

Silica coating on MNS is popular because of the ease of synthesis and aqueous stability. By hydrolyzing silica precursors in basic solution, a uniform and thickness-controllable silica coating on MNS was obtained [76, 94, 95]. The silica shell have been used as a carrier for anticancer drugs (e.g., paclitaxel) and fluorescent molecules (e.g., fluorescein isothiocyanate (FITC)) [96]. Imparting additional functionalities to the silica coating have enabled targeting and labeling functionality. By addition of 3-aminopropyl-triethoxysilane (APS) to the silica precursors, we have coated silica shells with primary amine groups with controlled thickness [97]. Similarly, by reacting APS with isothiocyanate functionalized fluorescent dyes, Lu et al. were able to develop fluorescent MNS, a multimodal diagnostic agent [94]. Currently, silica coated MNS is available as ferumoxil (AMI-121), an orally ingested T ₂ contrast agent for delineation of the intestinal loops from adjacent tissues and organs [21].

Passive targeting uses the predetermined physicochemical properties of MNS to specifically migrate to selected tissues. The most common example of passive targeting is the enhanced permeability and retention (EPR) effect where MNS smaller than 200 nm can accumulate in many tumor tissues passively in solid tumors [111]. The compromised vasculature of a solid tumor facilitates passive MNS extravasation from the circulation into the tumor interstitium (Fig. 7) [112]. By contrast, endothelial cells of normal tissue vessels are closely packed and present a barrier for MNS penetration. However, passive targeting is limited to specific tumors since success of EPR effect depends on a number of factors such as lymphatic drainage rate, degree of capillary disorder, and blood flow which varies in different tumor types [113, 114].

Because passive targeting is available for only certain types of tumors and does not necessarily guarantee internalization of MNS by targeted cells, MNS can be modified with tumor-selective agents to employ active targeting [115, 116]. These agents are complementary to unique receptors that are overexpressed or present on tumor cells. A variety of targeting agents have been used for MNS, depending on the specific target, and these are reviewed elsewhere [5, 10]. Some of the studies include: small organic molecules [115, 117, 118], peptides [119], proteins [120], and antibodies [27]. The density and molecular organization of these ligands significantly influence MNS binding to target cells due to the multivalency phenomenon [121]. Some of the targeting agents can be used to facilitate MNS internalization into cells, primarily via endocytosis [5]. However, synthesis of these targeting agents is expensive and involves complicated chemistry. Therefore, the process of scaling up the synthesis is challenging and may be a hurdle for clinical applications.

Accumulation of MNS can be realized by applying external magnetic field on the target site, a unique feature for MNS. Magnetic targeting has been studied for a number of tumor models [122, 123]. This technique was successfully implemented in a clinical trial to deliver the chemotherapeutic, doxorubicin, to hepatocarcinoma cells [124]. David et al. have explored magnetic targeting to brain tumors with PEI functionalized MNS [102, 122]. While successful, the efficacy of magnetic targeting is limited to target tissue that is close to the body’s surface, since the magnetic field strength decreases with the distance from the magnetic source.

According to Eq. 1, the R ₂ of MNS is proportional to saturation magnetization and volume. Since the saturation magnetization of an MNS is proportional to its size due to surface spin-canting effects, the r ₂ value of MNS can be increased by increasing the size of MNS. However, for biological imaging applications, the hydrodynamic size of MNS should be below 100 nm in order to have longer circulation times and to avoid nonspecific uptake [145]. Cheon et al. investigated the size effect where Fe₃O₄ MNS of diameter 4, 6, 9, and 12 nm resulted in the r ₂ relaxivity of 78, 106, 130, and 218 mM⁻¹ s⁻¹, respectively (Fig. 8). The MR contrast changed from light gray to black or from red to blue in color-coded images (Fig. 8) [27]. The 9 nm Fe₃O₄ MNS conjugated with Herceptin was used to image a breast cancer cell line, SK-BR-3, which possesses overexpressed HER2/neu cancer markers. Herceptin was selected due to its specific binding properties against a HER2/neu receptor. In the T ₂-weighted MR images, treatment of 9 nm Fe₃O₄ MNS-Herceptin probe conjugates to the SK-BR03 breast cancer cell lines resulted in the significant negative contrast of the MR images compared to nontreated cell lines [27]. Chen et al. studied the size effect of polyvinylpyrrolidone (PVP)-coated Fe₃O₄ MNS (core size 8, 23, 37, and 65) on MRI of hepatic lesions in vivo [146]. PVP-Fe₃O₄ MNS with core size 37 and 65 nm showed higher r ₂ relaxivity (239 and 248 mM⁻¹ s⁻¹, respectively) compared to other sizes. When administered in nude mice bearing orthotopic Huh7 liver cancer, PVP-Fe₃O₄ MNS with 37 nm core size showed higher contrast change compared to Feridex in T ₂ and T ₂* weighted MR images [146].