Magnetic Nanoparticle-Based High-Performance Positive and Negative Magnetic Resonance Imaging Contrast Agents

In recent decades, magnetic nanoparticles (MNPs) have attracted considerable research interest as versatile substances for various biomedical applications, particularly as contrast agents in magnetic resonance imaging (MRI). Depending on their composition and particle size, most MNPs are either paramagnetic or superparamagnetic. The unique, advanced magnetic properties of MNPs, such as appreciable paramagnetic or strong superparamagnetic moments at room temperature, along with their large surface area, easy surface functionalization, and the ability to offer stronger contrast enhancements in MRI, make them superior to molecular MRI contrast agents. As a result, MNPs are promising candidates for various diagnostic and therapeutic applications. They can function as either positive (T1) or negative (T2) MRI contrast agents, producing brighter or darker MR images, respectively. In addition, they can function as dual-modal T1 and T2 MRI contrast agents, producing either brighter or darker MR images, depending on the operational mode. It is essential that the MNPs are grafted with hydrophilic and biocompatible ligands to maintain their nontoxicity and colloidal stability in aqueous media. The colloidal stability of MNPs is critical in order to achieve a high-performance MRI function. Most of the MNP-based MRI contrast agents reported in the literature are still in the developmental stage. With continuous progress being made in the detailed scientific research on them, their use in clinical settings may be realized in the future. In this study, we present an overview of the recent developments in the various types of MNP-based MRI contrast agents and their in vivo applications.


Introduction
Magnetic resonance imaging (MRI) is a widely used, advanced and effective diagnostic imaging technique owing to its distinct advantages over other imaging modalities [1][2][3][4]. Its unique features include high spatial resolution [5,6], excellent soft-tissue contrast [7][8][9], lack of ionizing radiation risk [10,11], and high imaging depths. Owing to these factors, MRI is often the preferred choice for diagnosing various medical conditions, particularly in pregnant women and children who may be sensitive to ionizing radiation [10,12,13]. This is because MRI is operated under low-energy radiofrequency radiation in contrast to invasive X-ray computed tomography (CT) and positron emission tomography (PET) scans which use high-energy ionizing radiation and can damage cells and tissues, causing side effects. Although MRI supports high spatial resolution (0.05-0.5 mm) [5], its sensitivity is relatively low [5] because of the inherently small population difference between two proton spin energy states in nuclear magnetic resonance, making the detection of small biological events, such as early-stage small lesions, difficult. However, MRI sensitivity can Table 1. Classification of imaging modes of MNPs (++: a major application; +: a minor application).

Tb-Based MNP
T 1 + ++ ++ T 2 ++ + + ++ ++ ++ Dual-modal T 1 and T 2 + + + surface-coating ligand layer that ensures colloidal stability and nontoxicity. High colloidal stability in aqueous media is necessary for obtaining high-performance MRI function and maximal MR signal enhancement. Otherwise, the precipitation of MNPs can result in a reduced interaction between MNPs and water proton spins, resulting in reduced MR signal enhancement. In this review, we focused on newly developed, MNP-based MRI contrast agents and their in vivo applications. The review includes MNPs made of Gd, Mn, Fe, Dy, Ho, and Tb, which function as T1 MRI contrast agents; T2 MRI contrast agents; or dual-modal T1 and T2 MRI contrast agents.

The Principle of Imaging Mode (T1 or T2 or Both)
The effectiveness of MNPs as MRI contrast agents depends on their r1 and r2 values, and their r2/r1 ratios [46,47]. As the longitudinal (T1) proton spin relaxation always occurs with the transverse (T2) proton spin relaxation, but not the other way around, the ideal T1 MRI contrast agent should have high r1 values and r2/r1 ratios close to 1.0. In contrast, the ideal T2 MRI contrast agent should have high r2 values and high r2/r1 ratios to solely induce T2 proton spin relaxation [46,47]. Dual-modal T1 and T2 MRI contrast agents should have both high r1 and r2 values, with r2/r1 ratios between those of ideal T1 and ideal T2 MRI contrast agents. The primary parameter that affects r1 and r2 values and r2/r1 ratios is the MNP composition. The particle size and surface-coating ligand are also important. The types of imaging modes possible with MNPs are summarized in Table 1 and Scheme 1. Table 1. Classification of imaging modes of MNPs (++: a major application; +: a minor application).
Recent studies indicate that the particle diameter of Fe-based NPs is critical in achieving optimal T 1 proton spin relaxation because smaller NPs offer greater surface exposure of Fe 3+ to the surrounding water protons [52][53][54][55][56][57][58]. Park and co-workers reported that the optimal particle diameter range of Gd-based NPs for maximal T 1 proton spin relaxation or the maximal r 1 value is 1-2.5 nm [40]. They proposed the cooperative effects of surface Gd 3+ in accelerating T 1 proton spin relaxation to explain this.

T 2 Imaging Mode
MNPs become magnetized upon exposure to an external magnetic field, resulting in local magnetic field fluctuations around the MNPs because of their thermal motion in solution [59][60][61]. The local magnetic field fluctuations accelerate the transverse (T 2 ) proton spin relaxation (so-called outer sphere model [48][49][50]), which corresponds to the dephasing of proton spin magnetic moments along the xy plane. MNPs with high magnetic moments, such as SPIONs, can generate high local magnetic field fluctuations that induce strong T 2 proton spin relaxation, resulting in high r 2 values. Note that MNPs made of Dy, Ho, and Tb exhibit appreciable paramagnetic moments at room temperature and thus exhibit appreciable r 2 values [23][24][25][26][27][28][62][63][64]. They solely induce T 2 proton spin relaxation, resulting in negligible r 1 values. Notably, these paramagnetic moments increase with increasing MR field, resulting in even higher r 2 values than those of SPIONs at very high MR fields [62].

T 1 and T 2 Dual-Imaging Mode
Conventional MRI contrast agents can serve as either T 1 or T 2 MRI contrast agents. However, the introduction of T 1 and T 2 dual-modal MRI contrast agents has enabled more accurate diagnosis of disease via complementary T 1 and T 2 MR images [65][66][67][68][69][70][71]. Dualmodal MR images can be easily obtained by changing the operational mode of the same MRI scanner, unlike other dual-modal imaging techniques (such as MRI-CT, CT-PET, and MRI-PET) that require a combination of two different imaging machines that are expensive, inconvenient, and time-consuming.
Designing dual-modal MNPs requires careful synthetic strategies to ensure they function in a dual-mode. The composition, particle diameter, and surface-coating ligand of MNPs can be controlled to optimize the r 1 and r 2 values [65,[72][73][74][75][76]. For instance, small and hydrophilic surface-coating ligands can attract water molecules close to the MNPs, resulting in high r 1 and r 2 values, whereas large and hydrophobic surface-coating ligands do not, resulting in small r 1 and r 2 values [72,73,76]. Larger MNPs can provide higher r 2 values and higher r 2 /r 1 ratios owing to their enhanced magnetic moments and reduced amounts of surface metal ions and thus are more suitable as T 2 MRI contrast agents [74]. The composition of MNPs may be controlled to obtain r 1 and r 2 values suitable for dual-modal imaging [65]. Park et al. synthesized D-glucuronic acid-coated Gd 2 O 3 NPs (mean particle diameter = 1 nm) in polyol solvent [40]. The transmission electron microscope (TEM) image is presented in Figure 1a(i). The synthesized NPs exhibited a high r 1 value of 9.9 s −1 mM −1 and thus high positive T 1 contrasts in mouse brain tumors after intravenous injection at 1.5 T (Figure 1a(ii)).

Examples of MNP-Based MRI Contrast Agents
Recently, Yang et al. reported polyvinylpyrrolidone (PVP)-coated Gd 2 O 3 NPs (PVP Mw = 10,000 amu and mean particle diameter = 2.5 ± 0.8 nm) with a high r 1 value of 10.8 mM −1 s −1 at 3 T [77]. Appreciable contrasts in T 1 MR images of the tumor, kidney, bladder, and liver were observed after intravenous injection at 3 T.
Dai et al. reported a comparison study between polyethylene glycol (PEG)-coated (or PEGylated)-Gd 2 O 3 NPs (PEG Mw = 600 amu and mean hydrodynamic diameter = 36.35 nm) and commercially available Magnevist (or gadopentetic acid) [78]. The PEGylated-Gd 2 O 3 NPs exhibited a higher r 1 value of 29.0 mM −1 s −1 than that (= 4.2 mM −1 s −1 ) of Magnevist at 3 T. The T 1 MR images of tumor-bearing mice were obtained 3 T (Figure 1b). The contrast enhancement of PEGylated-Gd 2 O 3 NPs in the tumor was stronger than that of Magnevist at the same injection dose, confirming the superiority of PEGylated-Gd 2 O 3 NPs, compared with Magnevist.

Gd-Based NPs
Among MNPs, Gd-based NPs possess the most suitable relaxivity properties for T1 MRI contrast agents, owing to Gd 3+ having the highest spin magnetic moment (s = 7/2) among the elements in the periodic table. The r1 values of Gd-based NPs are higher than those [48,50] of Gd chelates because of the high density of Gd 3+ per NP, making them high-performance T1 MRI contrast agents. Among various types of Gd-based NPs, such as Gd2O3, GdF3, and NaGdF4 NPs, the Gd2O3 NPs have been most intensively investigated thus far.
Park et al. synthesized D-glucuronic acid-coated Gd2O3 NPs (mean particle diameter = 1 nm) in polyol solvent [40]. The transmission electron microscope (TEM) image is presented in Figure 1a(i). The synthesized NPs exhibited a high r1 value of 9.9 s -1 mM -1 and thus high positive T1 contrasts in mouse brain tumors after intravenous injection at 1.5 T (Figure 1a(ii)).
Dai et al. reported a comparison study between polyethylene glycol (PEG)-coated (or PEGylated)-Gd2O3 NPs (PEG Mw = 600 amu and mean hydrodynamic diameter = 36.35 nm) and commercially available Magnevist (or gadopentetic acid) [78]. The PEGylated-Gd2O3 NPs exhibited a higher r1 value of 29.0 mM −1 s −1 than that (= 4.2 mM −1 s −1 ) of Magnevist at 3 T. The T1 MR images of tumor-bearing mice were obtained 3 T ( Figure  1b). The contrast enhancement of PEGylated-Gd2O3 NPs in the tumor was stronger than that of Magnevist at the same injection dose, confirming the superiority of PEGylated-Gd2O3 NPs, compared with Magnevist.

Mn-Based NPs
Mn-based NPs have emerged as alternatives to Gd-based NPs because of their lower toxicity. Mn-based NPs have shown appreciable paramagnetic moments at room temperature owing to Mn 2+ (s = 5/2), but their paramagnetic moments are lower than those of Gd-based NPs because Mn 2+ (s = 5/2) has a lower spin magnetic moment than Gd 3+ (s = 7/2).

Mn-Based NPs
Mn-based NPs have emerged as alternatives to Gd-based NPs because of their lower toxicity. Mn-based NPs have shown appreciable paramagnetic moments at room temperature owing to Mn 2+ (s = 5/2), but their paramagnetic moments are lower than those of Gd-based NPs because Mn 2+ (s = 5/2) has a lower spin magnetic moment than Gd 3+ (s = 7/2).
Li et al. synthesized PEG-coated MnO NPs (PEG Mw = 600 amu and mean particle diameter = 1.9 nm), with a high r 1 value of 12.942 s −1 mM −1 and an r 2 /r 1 ratio of 4.66 at 3 T [80]. They conjugated the PEG-MnO NPs with AS1411 aptamer to target tumors, resulting in AS1411-PEG-MnO NPs that retained contrast in the tumor for up to 7 days after intravenous injection. However, the contrast provided by PEG-MnO NPs disappeared after 1 day of injection because of their lack of tumor targeting.
Xiao et al. prepared ligand-free Mn 3 O 4 NPs (mean particle diameter = 9 nm) using laser ablation in aqueous media [81], as shown in a TEM image presented in Figure 2b(i). These uncoated Mn 3 O 4 NPs had an r 1 value of 8.26 mM −1 s −1 at 3 T, which did not significantly affect the cell viability, owing to their low toxicity. T 1 MR images at 3 T revealed high positive contrast enhancements in the xenografted tumor 30 min after intravenous injection (Figure 2b(ii)), demonstrating their potential as a T 1 MRI contrast agent.
NPs with proper ligands such as ZDS, not DMSA, is crucial for the penetration of the NPs through the brain-blood barrier.
Li et al. synthesized PEG-coated MnO NPs (PEG Mw = 600 amu and mean particle diameter = 1.9 nm), with a high r1 value of 12.942 s −1 mM −1 and an r2/r1 ratio of 4.66 at 3 T [80]. They conjugated the PEG-MnO NPs with AS1411 aptamer to target tumors, resulting in AS1411-PEG-MnO NPs that retained contrast in the tumor for up to 7 days after intravenous injection. However, the contrast provided by PEG-MnO NPs disappeared after 1 day of injection because of their lack of tumor targeting.
Xiao et al. prepared ligand-free Mn3O4 NPs (mean particle diameter = 9 nm) using laser ablation in aqueous media [81], as shown in a TEM image presented in Figure 2b(i). These uncoated Mn3O4 NPs had an r1 value of 8.26 mM −1 s −1 at 3 T, which did not significantly affect the cell viability, owing to their low toxicity. T1 MR images at 3 T revealed high positive contrast enhancements in the xenografted tumor 30 min after intravenous injection (Figure 2b(ii)), demonstrating their potential as a T1 MRI contrast agent.

Fe-Based NPs
Because of the high magnetic moments and better biocompatibility of Fe-based NPs compared with other metal-based NPs, they have been extensively investigated as T2 MRI contrast agents.
Wang et al. reported the one-pot synthesis of ultrasmall SPIONs (uBSPIOs) using bovine serum albumin as a scaffold [83]. The resulting uBSPIOs (mean particle diameter = 4.78 ± 0.55 nm) exhibited a high r2 value of 444.56 ± 8.82 mM −1 s −1 at 7 T. These uBSPIOs provided high negative contrasts in the T2 MR images of the tumor. In addition, the uB-SPIOs demonstrated no cytotoxicity in vitro and negligible organ toxicity in vivo.

Fe-Based NPs
Because of the high magnetic moments and better biocompatibility of Fe-based NPs compared with other metal-based NPs, they have been extensively investigated as T 2 MRI contrast agents.
Wang et al. reported the one-pot synthesis of ultrasmall SPIONs (uBSPIOs) using bovine serum albumin as a scaffold [83]. The resulting uBSPIOs (mean particle diameter = 4.78 ± 0.55 nm) exhibited a high r 2 value of 444.56 ± 8.82 mM −1 s −1 at 7 T. These uBSPIOs provided high negative contrasts in the T 2 MR images of the tumor. In addition, the uBSPIOs demonstrated no cytotoxicity in vitro and negligible organ toxicity in vivo.
Leal et al. synthesized PEGylated SPIONs (mean particle diameter = 6 nm) with PEG molecular weight (MW) ranging from 600 to 8000 amu [84]. The PEG3000-SPIONs (PEG Mw = 3000 amu) exhibited better in vivo performance, with longer circulation times and slower liver uptake than other MW PEG-coated SPIONs. This was because high MW PEGs (6000-8000 amu) led to NP aggregations, whereas low MW PEGs (≤1500 amu) could not stabilize the NPs in physiological media. The PEG1500-, PEG3000-, PEG6000-and PEG8000-SPIONs had r 2 values of 90,103,190, and 180 mM −1 s −1 at 1.5 T and the r 2 values of 114, 151, 253, and 254 mM −1 s −1 at 9.4 T, respectively. After intravenous injection into mice tails at 9.4 T, the PEGylated SPIONs exhibited negative contrast enhancements in T 2 MR images, demonstrating their potential as T 2 MRI contrast agents.
Lee et al. synthesized ferrimagnetic iron oxide nanocubes (FIONs) (mean edge length = 22 nm, as presented in a TEM image in Figure 3b(i)) and encapsulated them in PEGphospholipids to obtain water-dispersible FIONs (WFIONs) [85]. The WFIONs exhibited an extremely high r 2 value of 761 mM −1 s −1 at 3 T, which is close to the theoretically predicted maximum r 2 value of~800 mM −1 s −1 for 22 nm sized NPs with a saturation magnetization of 106 emu/g. Notably, their study demonstrated that FION aggregates could generate a very strong magnetic field that caused nearby water proton spins to completely lose their phase, leading to a decrease in their r 2 value as the NP size increased. In vivo MR images revealed a distinct signal decrease in the tumor site 1 h after intravenous injection (Figure 3b(ii)).
Maurea et al. reported the clinical applications of a commercial T 2 MRI contrast agent Resovist in humans [38]. Resovist is carboxydextrane-coated SPION with a hydrodynamic diameter ranging from 45 to 60 nm and r 2 and r 1 values of 151.0 and 25.4 mM −1 s −1 , respectively [35]. As shown in Figure 3c, hepatocellular carcinoma could be more clearly observed after intravenous injection compared with no injection (labeled as before) [38].
into mice tails at 9.4 T, the PEGylated SPIONs exhibited negative contrast enhancements in T2 MR images, demonstrating their potential as T2 MRI contrast agents.
Lee et al. synthesized ferrimagnetic iron oxide nanocubes (FIONs) (mean edge length = 22 nm, as presented in a TEM image in Figure 3b(i)) and encapsulated them in PEG-phospholipids to obtain water-dispersible FIONs (WFIONs) [85]. The WFIONs exhibited an extremely high r2 value of 761 mM −1 s −1 at 3 T, which is close to the theoretically predicted maximum r2 value of ~800 mM −1 s −1 for 22 nm sized NPs with a saturation magnetization of 106 emu/g. Notably, their study demonstrated that FION aggregates could generate a very strong magnetic field that caused nearby water proton spins to completely lose their phase, leading to a decrease in their r2 value as the NP size increased. In vivo MR images revealed a distinct signal decrease in the tumor site 1 h after intravenous injection (Figure 3b(ii)).
Maurea et al. reported the clinical applications of a commercial T2 MRI contrast agent Resovist in humans [38]. Resovist is carboxydextrane-coated SPION with a hydrodynamic diameter ranging from 45 to 60 nm and r2 and r1 values of 151.0 and 25.4 mM -1 s -1 , respectively [35]. As shown in Figure 3c, hepatocellular carcinoma could be more clearly observed after intravenous injection compared with no injection (labeled as before) [38].

Dy-Based NPs
Dy-based NPs have gained significant interest as a new class of T 2 MRI contrast agents because of their appreciable paramagnetic moments at room temperature. Dy has the highest effective magnetic moment (µ eff ) (10.65 µ B ) among the elements in the periodic table [51]. Currently, most MRI contrast agents are operated at clinical MR fields (1.5-3 T). However, Dy-based NPs can become more powerful at higher MR fields such as 7 and 9.4 T because their paramagnetic moment increases with the increasing MR field. This will trigger a new opportunity to generate powerful T 2 MRI contrast agents suitable for high-field MRI scanners. In addition to T 2 relaxation by outer sphere model, Curie-spin relaxation is also important for Dy-based NPs [86,87]. This is because T 2 relaxation is induced by the modulated magnetic dipolar interaction between proton spins and thermally averaged electronic spins (or Curie-spins) of Dy 3+ in NPs by NP motion. This Curie-spin relaxation becomes important in high MR fields.
González-Mancebo et al. synthesized rhombus-like DyF 3 NPs (mean length × width = 110 × 50 nm) in ethylene glycol, which served as the solvent and surface-coating ligand [88]. The DyF 3 NPs exhibited a remarkably high r 2 value of 380.4 mM −1 s −1 at 9.4 T, with a high r 2 /r 1 ratio of 559.37, demonstrating their potential as highly effective T 2 MRI contrast agents.
Kattel et al. reported on the effectiveness of D-glucuronic acid-coated Dy 2 O 3 NPs and Dy(OH) 3 nanorods as T 2 MRI contrast agents [89]. The D-glucuronic acid-coated Dy 2 O 3 NPs (mean particle diameter = 3.2 nm) and Dy(OH) 3 nanorods (mean diameter × length = 20 × 300 nm) exhibited negligible r 1 values and high r 2 values of 65.04 and 181.57 s −1 mM −1 at 1.5 T, respectively. Dy 2 O 3 NPs and Dy(OH) 3 nanorods also produced clear negative contrast enhancements of T 2 MR images of mouse liver and kidneys at 3 T after intravenous injection.
Recently, Yue et al. synthesized hydrophilic, nearly nontoxic, amorphous carboncoated Dy 2 O 3 NPs (mean particle diameter = 3.0 nm, as presented in a TEM image in Figure 4a(i)) [90]. The carbon coating was achieved through dextrose polymerization on the Dy 2 O 3 NP surfaces, which left hydroxyl groups on the NP surfaces and made the amorphous carbon-coated Dy 2 O 3 NPs colloidally stable in aqueous media and nearly nontoxic. The amorphous carbon-coated NPs had r 1 and r 2 values of 0.1 and 5.7 s −1 mM −1 at 3 T, respectively, with an r 2 /r 1 ratio of 57. In vivo T 2 MR images of mouse kidneys at 3 T exhibited negative contrasts after intravenous injection, demonstrating their potential as T 2 MRI contrast agents (Figure 4a(ii)). In addition, the NPs exhibited broad photofluorescence at 490 nm (400-600 nm) upon excitation at 370 nm due to the fluorescent nature of the amorphous carbon-coating layers on the NP surfaces. Therefore, amorphous carbon-coated Dy 2 O 3 NPs are suitable as dual-modal T 2 MRI-fluorescence imaging (FI) agents.
Marasini et al. developed colloidally stable poly(acrylic) acid (PAA)-coated Dy 2 O 3 NPs (PAA Mw = 1800 amu and mean particle diameter = 1.7 nm, as presented in a TEM image in Figure 4b(i)) using a simple polyol synthesis [91]. The r 1 value was negligible, but the r 2 value increased with increasing MR field such that it was 2.01 s −1 mM −1 at 3 T and 11.31 s −1 mM −1 at 9.4 T. In vivo T 2 MR images at 3 T exhibited clear negative contrast enhancements in mouse livers after intravenous injection (Figure 4b(ii)), demonstrating their potential as a T 2 MRI contrast agent.

Ho-Based NPs
Ho-based NPs exhibit appreciable paramagnetic moments at room temperature, similar to Dy-based NPs, owing to Ho with high μeff of 10.60 μB, the second highest value among elements in the periodic table [51]. Consequently, they are potential candidates for T2 MRI contrast agents at high MR fields similar to Dy-based NPs.

Ho-Based NPs
Ho-based NPs exhibit appreciable paramagnetic moments at room temperature, similar to Dy-based NPs, owing to Ho with high µ eff of 10.60 µ B, the second highest value among elements in the periodic table [51]. Consequently, they are potential candidates for T 2 MRI contrast agents at high MR fields similar to Dy-based NPs.
Gómez-González et al. synthesized cube-shaped HoPO 4 NPs that were grafted with PAA (Mw = 1800 amu) to investigate the relationship between r 2 value and MR field (1.44 and 9.4 T) and between r 2 value and NP diameter (27,48, and 80 nm) [93]. The r 2 value of HoPO 4 @PAA NPs increased with increasing NP diameter at 1.44 T, but failed for the largest NPs at 9.4 T because of their aggregation; consequently, the 48 nm NPs exhibited the highest r 2 value of 489.91 mM −1 s −1 at 9.4 T. In vivo studies using 48 nm HoPO 4 @PAA NPs at 9.4 T revealed that, after intravenous injection, the NPs showed distinct T 2 contrasts in both the liver and spleen, demonstrating their potential as T 2 MRI contrast agents.
Atabaev et al. reported an r 2 value of 23.47 mM −1 s −1 at 1.5 T for PEG-coated Ho 2 O 3 NPs (PEG Mn = 4000 amu) with a particle diameter of 67-81 nm [94], which is high enough to be used as T 2 MRI contrast agents.
González-Mancebo et al. investigated the r 2 values of HoF 3 NPs at 9.4 T as a function of particle size and composition [88]. They synthesized two different types of HoF 3 NPs in ethylene glycol, which served as the solvent and surface-coating ligand; ellipsoid-like HoF 3 NPs (HoF-el) (mean length × width = 70 × 30 nm) and rhombus-like HoF 3 NPs (HoF-rh) (mean length × width = 110 × 50 nm). The NPs exhibited r 2 values of 349.98 mM −1 s −1 and 608.4 mM −1 s −1 at 9.4 T for HoF-el and HoF-rh, respectively. The increase in r 2 value with increasing particle size was attributed to the higher magnetization of larger NPs than smaller ones. The high r 2 values indicated that the synthesized NPs should be useful as high-field T 2 MRI contrast agents.
Recently, Zhang et al. conducted a study on PEG-HoF 3 NPs (PEG Mw = 4000 amu and mean particle diameter = 38 nm) as a T 2 MRI contrast agent for cancer diagnosis [95]. The NPs had an r 2 value of 117.51 mM −1 s −1 at 7 T, and 24 h after intravenous injection, negative (or darker) contrasts in T 2 MR images in the tumors of tumor-bearing mice were observed because of the accumulation of NPs in the tumor, demonstrating the potential of the NPs as a T 2 MRI contrast agent.
Marasini et al. synthesized PAA-coated Ho 2 O 3 NPs (PAA Mw = 1800 amu and mean particle diameter = 1.7 nm) using one-pot polyol synthesis [96]. The TEM image is presented in Figure 5a(i). The NPs were nearly nontoxic and colloidally stable in aqueous media without precipitation after synthesis because of their hydrophilic and biocompatible PAA coating. The PAA-coated Ho 2 O 3 NPs exhibited an appreciable r 2 value of 1.44 s −1 mM −1 at 3 T and an enhanced r 2 value of 9.20 s −1 mM −1 at 9.4 T. In vivo T 2 MR images of the liver and kidneys exhibited strong negative contrast enhancements at 3 T, and stronger negative contrast enhancements at 9.4 T, confirming the effectiveness of the NPs as T 2 MRI contrast agents at high-MR fields (Figure 5a(ii)).

Tb-Based NPs
Tb-based NPs have shown great potential as T2 MRI contrast agents at high MR fields owing to their appreciable paramagnetic moments at room temperature, which were similar to Dy-and Ho-based NPs because Tb has μeff of 9.72 μB [51]. Despite their promising potential as T2 MRI contrast agents at high MR fields, only a few studies on Tb-based NPs have been reported.
Zheng et al. synthesized PEI-coated TbF3 NPs (PEI Mn = 25,000 amu) using a facile solvothermal method [99]. The NPs had a plate morphology (mean particle diameter  thickness = 160  29 nm), as presented in the TEM image in Figure 6a(i). They obtained high r2 values of 6.54 and 395.77 mM -1 s -1 at 0.5 and 7 T, respectively. An in vivo T2 MRI study at 7 T revealed that after injection into mouse tail veins, the MR signal intensities in the liver, spleen, and kidneys decreased significantly after injection, indicating the accumulation of NPs in these organs. An example of a liver MR image 15 min after injection is presented in Figure 6a(ii).

Tb-Based NPs
Tb-based NPs have shown great potential as T 2 MRI contrast agents at high MR fields owing to their appreciable paramagnetic moments at room temperature, which were similar to Dy-and Ho-based NPs because Tb has µ eff of 9.72 µ B [51]. Despite their promising potential as T 2 MRI contrast agents at high MR fields, only a few studies on Tb-based NPs have been reported.
Zheng et al. synthesized PEI-coated TbF 3 NPs (PEI Mn = 25,000 amu) using a facile solvothermal method [99]. The NPs had a plate morphology (mean particle diameter × thickness = 160 × 29 nm), as presented in the TEM image in Figure 6a(i). They obtained high r 2 values of 6.54 and 395.77 mM −1 s −1 at 0.5 and 7 T, respectively. An in vivo T 2 MRI study at 7 T revealed that after injection into mouse tail veins, the MR signal intensities in the liver, spleen, and kidneys decreased significantly after injection, indicating the accumulation of NPs in these organs. An example of a liver MR image 15 min after injection is presented in Figure 6a(ii).

Tb-Based NPs
Tb-based NPs have shown great potential as T2 MRI contrast agents at high MR fields owing to their appreciable paramagnetic moments at room temperature, which were similar to Dy-and Ho-based NPs because Tb has μeff of 9.72 μB [51]. Despite their promising potential as T2 MRI contrast agents at high MR fields, only a few studies on Tb-based NPs have been reported.
Zheng et al. synthesized PEI-coated TbF3 NPs (PEI Mn = 25,000 amu) using a facile solvothermal method [99]. The NPs had a plate morphology (mean particle diameter  thickness = 160  29 nm), as presented in the TEM image in Figure 6a(i). They obtained high r2 values of 6.54 and 395.77 mM -1 s -1 at 0.5 and 7 T, respectively. An in vivo T2 MRI study at 7 T revealed that after injection into mouse tail veins, the MR signal intensities in the liver, spleen, and kidneys decreased significantly after injection, indicating the accumulation of NPs in these organs. An example of a liver MR image 15 min after injection is presented in Figure 6a(ii). Marasini et al. synthesized D-glucuronic acid-coated Tb 2 O 3 NPs (mean particle diameter = 2.0 nm) as a potential dual-modal T 2 MRI-FI agent because Tb also emits photons in the 545 nm region [100]. The D-glucuronic acid-coated NPs exhibited r 2 values of 7.68 mM −1 s −1 at 1.5 T, 33.97 mM −1 s −1 at 3 T, and 53.67 mM −1 s −1 at 9.4 T, indicating that they have r 2 values suitable as T 2 MRI contrast agents. Furthermore, the NPs exhibited fluorescence in the green region, making them suitable as dual-modal T 2 MRI-FI agents.
Recently, Caro et al. investigated the potential of PEG-coated Tb-based nanorods (PEG-TbNRs) as multimodal bioimaging agents [24]. The PEG-TbNRs (PEG Mw = 3000 amu and mean particle diameter × length = 2 × 9 nm) exhibited high colloidal stability and excellent luminescent, magnetic, and X-ray attenuation properties. The scanning TEM (STEM) image demonstrating the nanorod morphology is presented in Figure 6b(i). The r 2 values of PEG−TbNRs at 1.44 and 9.4 T were estimated to be 10.4 and 48.5 mM −1 s −1 , respectively. In vivo T 2 MR images at 9.4 T exhibited appreciable negative contrast enhancements in the liver and kidneys after intravenous injection (Figure 6b(ii)), demonstrating the potential of PEG-TbNRs as T 2 MRI contrast agents.
Miao et al. synthesized PAA-coated Fe 3 O 4 NPs (PAA = 1000 amu and mean particle diameter = 5.1 nm). The NPs exhibited r 1 and r 2 values of 10.52 and 38.97 mM −1 s −1 (r 2 /r 1 = 3.70) at 1.41 T, respectively [109]. The TEM image is presented in Figure 7a(i). The performance of the NPs as a dual-modal T 1 and T 2 MRI contrast agent was demonstrated from T 1 and T 2 MR images at 3 T, where positive contrasts were clearly observed in the rabbit vasculature (Figure 7a(ii)) and negative contrasts were clearly observed in the rabbit popliteal lymph node (dotted circle) (Figure 7a(iii)).
Li et al. developed monodispersed water-soluble and biocompatible ultrasmall magnetic iron oxide nanoparticles (mean particle diameter = 3.3 ± 0.5 nm) grafted with poly(methacrylic acid) (PMAC, M n = 6359 amu) in aqueous media using a high-temperature coprecipitation method [104]. The PMAC-grafted NPs exhibited r 1 = 8.3 and r 2 = 35.1 s −1 mM −1 (r 2 /r 1 = 4.2) at 4.7 T, and demonstrated their potential as dual-modal T 1 and T 2 MRI contrast agents. After intravenous injection, positive and negative contrasts were observed in the T 1 and T 2 MR images of mice liver and kidneys, respectively.
Recently, Marasini et al. developed a dual-modal T 1 and T 2 MRI contrast agent by coating Gd 2 O 3 NPs (mean particle diameter = 2 nm) with polyaspartic acid (PASP) (Mw =~9900 amu) using the one-pot polyol method [111]. The TEM image is presented in Figure 7b(i). The synthesized NPs exhibited high r 1 and r 2 values of 19.1 and 53.7 mM −1 s −1 (r 2 /r 1 = 2.8) at 3 T, respectively. After intravenous injection of PASP-coated Gd 2 O 3 NPs into the mice tails, T 1 and T 2 contrasts were observed in the T 1 and T 2 MR images of the mouse livers at 3 T, respectively (Figure 7b(ii)). This result showed that dual-modal T 1 and T 2 MRI contrast agents prepared using Gd 2 O 3 NPs could be obtained by choosing the appropriate hydrophilic polymers as surface-coating ligands, such as the PASP used in this study.
Another promising candidate for dual-modal T 1 and T 2 MRI contrast agents is Mnbased NPs. Niu et al. developed manganese oxide nanocluster-loaded (diameter < 2 nm) dual-mesoporous silica spheres (Mn-DMSS; diameter = 100-200 nm, as presented in the TEM image in Figure 7c(i)) [112]. Mn-DMSSs exhibited a high r 1 value of 10.1 mM −1 s −1 and a high r 2 value of 169.7 mM −1 s −1 (r 2 /r 1 = 16.8) at 3 T. An in vivo experiment on rats at 3 T demonstrated that Mn-DMSSs exhibited a 29% signal enhancement in the liver under T 1 imaging mode and a 28% signal decrease under T 2 imaging mode (Figure 7c(ii)), demonstrating the potential of Mn-DMSSs as dual-modal T 1 and T 2 MRI contrast agents. ~9900 amu) using the one-pot polyol method [111]. The TEM image is presented in Figure  7b(i). The synthesized NPs exhibited high r1 and r2 values of 19.1 and 53.7 mM −1 s −1 (r2/r1 = 2.8) at 3 T, respectively. After intravenous injection of PASP-coated Gd2O3 NPs into the mice tails, T1 and T2 contrasts were observed in the T1 and T2 MR images of the mouse livers at 3 T, respectively (Figure 7b(ii)). This result showed that dual-modal T1 and T2 MRI contrast agents prepared using Gd2O3 NPs could be obtained by choosing the appropriate hydrophilic polymers as surface-coating ligands, such as the PASP used in this study.
Another promising candidate for dual-modal T1 and T2 MRI contrast agents is Mn-based NPs. Niu et al. developed manganese oxide nanocluster-loaded (diameter < 2 nm) dual-mesoporous silica spheres (Mn-DMSS; diameter = 100-200 nm, as presented in the TEM image in Figure 7c(i)) [112]. Mn-DMSSs exhibited a high r1 value of 10.1 mM −1 s −1 and a high r2 value of 169.7 mM −1 s −1 (r2/r1 = 16.8) at 3 T. An in vivo experiment on rats at 3 T demonstrated that Mn-DMSSs exhibited a 29% signal enhancement in the liver under T1 imaging mode and a 28% signal decrease under T2 imaging mode (Figure 7c(ii)), demonstrating the potential of Mn-DMSSs as dual-modal T1 and T2 MRI contrast agents.

Colloidal Stability, Biocompatibility, and Renal Excretion
It is essential that MNPs are coated with hydrophilic and biocompatible ligands to maintain their nontoxicity and colloidal stability in aqueous media. The colloidal stability of MNPs is critical for achieving high-performance MRI function because precipitated NPs lessen or negligibly contribute to inducing proton spin relaxation.
Compared with MNPs made of Gd, Mn, Dy, Ho, and Tb, Fe-based NPs are more biocompatible because iron is consumed in the human body as an essential element [33]; for example, it is used in hemoglobin for oxygen binding. For this reason, several Fe-based MNPs as MRI contrast agents, such as Feridex, Sinerem, and Resovist, were commercialized with approval by the FDA, USA [34−38], over other metal-based MNPs.

Colloidal Stability, Biocompatibility, and Renal Excretion
It is essential that MNPs are coated with hydrophilic and biocompatible ligands to maintain their nontoxicity and colloidal stability in aqueous media. The colloidal stability of MNPs is critical for achieving high-performance MRI function because precipitated NPs lessen or negligibly contribute to inducing proton spin relaxation.
Compared with MNPs made of Gd, Mn, Dy, Ho, and Tb, Fe-based NPs are more biocompatible because iron is consumed in the human body as an essential element [33]; for example, it is used in hemoglobin for oxygen binding. For this reason, several Fe-based MNPs as MRI contrast agents, such as Feridex, Sinerem, and Resovist, were commercialized with approval by the FDA, USA [34][35][36][37][38], over other metal-based MNPs.
It is critical that MNPs should be nontoxic for biomedical applications [113,114]. Because MRI contrast agents are generally intravenously injected, it is preferred that MNPs are excreted through the renal system rather than the hepatobiliary pathway because the hepatobiliary excretion is relatively slow and MNPs could decompose during the excretion process, which would be toxic to the body. For example, free Gd 3+ ions liberated from Gd chelates into the body could cause NSF [22,29,30]. For renal excretion, MNPs should be ultrasmall with hydrodynamic diameters less than 5 nm [115][116][117] because the glomerular filtration diameter in the kidneys is 4.5-5 nm [115]. It is also essential that the kinetic stability of MNPs is high so that they do not decompose until they are excreted through the renal system as urine.
As summarized in Figure 8, MNPs as MRI contrast agents for safe, in vivo applications should be kinetically stable (i.e., no decomposition), coated with hydrophilic and biocompatible polymers for nontoxicity and colloidal stability, and ultrasmall with hydrodynamic diameters less than 5 nm for renal excretion. Under these conditions, MNPs can serve as high-performance MRI contrast agents which are superior to commercial molecular MRI contrast agents. In particular, the magnetic properties of MNPs made of Gd, Dy, Ho, and Tb are nearly size independent and thus can be made ultrasmall for renal excretion. They can strongly induce proton spin relaxation at high MR fields, implying that those MNPs are potential candidates for a new type of MRI contrast agents for high-field MRI scanners.
Because MRI contrast agents are generally intravenously injected, it is preferred that MNPs are excreted through the renal system rather than the hepatobiliary pathway because the hepatobiliary excretion is relatively slow and MNPs could decompose during the excretion process, which would be toxic to the body. For example, free Gd 3+ ions liberated from Gd chelates into the body could cause NSF [22,29,30]. For renal excretion, MNPs should be ultrasmall with hydrodynamic diameters less than 5 nm [115][116][117] because the glomerular filtration diameter in the kidneys is 4.5-5 nm [115]. It is also essential that the kinetic stability of MNPs is high so that they do not decompose until they are excreted through the renal system as urine.
As summarized in Figure 8, MNPs as MRI contrast agents for safe, in vivo applications should be kinetically stable (i.e., no decomposition), coated with hydrophilic and biocompatible polymers for nontoxicity and colloidal stability, and ultrasmall with hydrodynamic diameters less than 5 nm for renal excretion. Under these conditions, MNPs can serve as high-performance MRI contrast agents which are superior to commercial molecular MRI contrast agents. In particular, the magnetic properties of MNPs made of Gd, Dy, Ho, and Tb are nearly size independent and thus can be made ultrasmall for renal excretion. They can strongly induce proton spin relaxation at high MR fields, implying that those MNPs are potential candidates for a new type of MRI contrast agents for high-field MRI scanners.

Conclusions
MRI has emerged as a promising imaging modality for accurate disease diagnosis. Although the population difference between the two proton spin energy states is small because of a small energy difference between them, the high content of hydrogens from water (~60 wt.% of the human body) and other sources in the human body results in a large number of hydrogen protons. This plays a critical role in producing MR signals and contrasts.
The MR signals and contrasts can be improved by accelerating the proton spin relaxation with MRI contrast agents. Commercial molecular MRI contrast agents such as Gd chelates typically have low r1 and r2 values and short blood circulation times, thus necessitating large numbers of injection doses to achieve their detection level. However, MNP-based MRI contrast agents can provide enhanced MR signals and contrast com-

Conclusions
MRI has emerged as a promising imaging modality for accurate disease diagnosis. Although the population difference between the two proton spin energy states is small because of a small energy difference between them, the high content of hydrogens from water (~60 wt.% of the human body) and other sources in the human body results in a large number of hydrogen protons. This plays a critical role in producing MR signals and contrasts.
The MR signals and contrasts can be improved by accelerating the proton spin relaxation with MRI contrast agents. Commercial molecular MRI contrast agents such as Gd chelates typically have low r 1 and r 2 values and short blood circulation times, thus necessitating large numbers of injection doses to achieve their detection level. However, MNP-based MRI contrast agents can provide enhanced MR signals and contrast compared with those of molecular MRI contrast agents owing to their enhanced magnetic moments, high density of metals per NP, and longer blood circulation times. Moreover, MNP-based MRI contrast agents have great potential for disease therapy through drug delivery and targeting ligand conjugation on MNP surfaces. In addition, T 1 and T 2 dual-modal-imaging MNPs show great potential for improving disease diagnosis via complementary T 1 and T 2 MR images, which molecular MRI contrast agents cannot provide.
Among the MNPs, only iron oxide NPs were commercialized as T 2 MRI contrast agents, but most of them have been withdrawn from the market. Only Resovist is currently available in a few countries. Commercial T 1 MRI contrast agents are Gd chelates and currently hold the market for all MRI contrast agents. However, they have low sensitivity owing to their low relaxivity values and short imaging times owing to their short blood circulation times. Therefore, the development of new MRI contrast agents to overcome such shortcomings is desirable. MNPs may be the breakthrough because they have high r 1 and r 2 values and long blood circulation times.

Perspective
This review provides an overview of MNP-based MRI contrast agents composed of Gd, Mn, Fe, Dy, Ho, and Tb. Their high-performance as MRI contrast agents was highlighted via in vivo MRI studies. However, most of MNP-based MRI contrast agents reported in the literature are still in the development stage, with limited research carried out on in vitro and early-stage in vivo small animal studies. To increase the possibility of their safe use as high-performance MRI contrast agents in clinical trials in the future, several key issues such as toxicological effects, long-term stability, and pharmacokinetics must be addressed.
Tailoring of the particle size, morphology, composition, and surface-coating ligand is essential in achieving high-performance MNP-based MRI contrast agents. Multidisciplinary collaborative research can help advance the synthetic techniques and gain an understanding of the correlation between the fundamental physicochemical properties of MNPs and their biological behaviors in vivo and in vitro. With continuous progress in detailed scientific research on MNP-based MRI contrast agents, their use in clinical settings may become feasible in the future.