Rational Design of Magnetic Nanoparticles as T1–T2 Dual-Mode MRI Contrast Agents

Magnetic nanoparticles (MNPs), either paramagnetic or superparamagnetic depending on their composition and size, have been thoroughly studied as magnetic resonance imaging (MRI) contrast agents using in vitro and in vivo biomedical preclinical studies, while some are clinically used. Their magnetic properties responsible in some cases for high magnetization values, together with large surface area-to-volume ratios and the possibility of surface functionalization, have been used in MRI-based diagnostic and theranostics applications. MNPs are usually used as positive (T1) or negative (T2) MRI contrast agents, causing brightening or darkening of selected regions in MRI images, respectively. This review focusses on recent developments and optimization of MNPs containing Gd, Mn, Fe and other lanthanide ions which may function as dual-mode T1–T2 MRI contrast agents (DMCAs). They induce positive or negative contrast in the same MRI scanner upon changing its operational mode between T1-weighted and T2-weighted pulse sequences. The type of contrast they induce depends critically on their r2/r1 relaxivity ratio, which for DMCAs should be in the 2–10 range of values. After briefly discussing the basic principles of paramagnetic relaxation in MNPs, in this review, the basic strategies for the rational design of DMCAs are presented and typical examples are discussed, including in vivo preclinical applications: (1) the use of NPs with a single type of contrast material, Gd- or Mn-based NPs or superparamagnetic NPs with appropriate size and magnetization to provide T2 and T1 contrast; and (2) inclusion of both types of T1 and T2 contrast materials in the same nanoplatform by changing their relative positions.


Introduction
Magnetic resonance imaging (MRI) is one of the most prominent clinical imaging modalities resulting from its many favorable characteristics.These include non-invasiveness, use of low-energy radiofrequency radiation, in contrast to invasive X-ray computed tomography (CT), positron-emission tomography (PET) and single-photon emission computed tomography (SPECT), which use damaging high-energy ionizing radiation, providing tomographic images with large penetration depth of any area of the body, in contrast to optical imaging (OI), outstanding spatial resolution (50-100 µm) and remarkable soft tissue contrast (see Table 1 for a comparison of some properties of imaging modalities for clinical applications).The contrast in the MRI images is generated by differences in intensities of 1 H NMR resonances, most frequently of water protons, with an important contribution of lipids in some cases.These intensities are governed by several parameters, including the local tissue 1 H concentrations, diffusion and flow of water molecules, and most importantly by their intrinsic differences in proton spin-lattice or longitudinal relaxation times (T 1 ) and spin-spin or transverse relaxation times (T 2 ) [1,2].However, the sensitivity of MRI is relatively low, due to the small population difference between the two proton spin energy states in the presence of a magnetic field, making it difficult to detect small lesions and limiting the time resolution of the technique.In recent decades, there has been a trend toward clinical MRI equipment operating at higher magnetic field strengths (B 0 ), which improved its contrast-to-noise ratios and spatial resolution.Presently, medical diagnostic MRI is typically performed at B 0 = 0.5-3 T, and research equipment can operate at B 0 values up to 11 T.
MRI contrast agents (CAs) are often employed to enhance the contrast between normal and disease tissues by significantly decreasing their T 1 and T 2 values [3][4][5][6][7].This acceleration of the spin relaxation of water protons, in the region where they accumulate relative to surrounding tissues, results from oscillating magnetic fields produced by metal ions present in the CAs, allowing sensitive MRI detection of that region.Their efficacy is measured by their relaxivity, r 1 or r 2 , defined as the paramagnetic enhancement of the water proton relaxation rates, normalized to a 1 mM metal ion concentration [3][4][5].
Table 1.Comparison of some typical properties of imaging modalities for clinical applications [8,9].A class of MRI CAs contains strongly paramagnetic metal ions, such as Gd 3+ , Mn 2+ and Fe 3+ , encapsulated by a strongly binding chelating ligands, or in the form of paramagnetic nanoparticles (MNPs).These CAs increase the longitudinal relaxation rates of water protons significantly more than their transverse relaxation rates, giving rise to bright spots in T 1 -weighted (T 1w ) MRI images and are called positive or T 1 CAs [3][4][5][6][7]10].Another class of CAs, containing paramagnetic metal ions, such as Tb 3+ , Dy 3+ and Ho 3+ , in the form of paramagnetic chelates or NPs, or as superparamagnetic NPs, such as iron-oxide NPs (SPIONs), induce protons in their vicinity to preferably undergo spin-spin relaxation, originating negative (or dark) contrast in T 2 -weighted MRI images and are called negative or T 2 CAs.In addition to r 1 and r 2 , the r 2 /r 1 ratio (always ≥1, as T 2 ≤ T 1 ) represents another important factor in identifying the class of MRI CAs.This ratio should be close to one for a T 1 CA (positive contrast), while it is large for a T 2 CA [10][11][12][13].

Technique
Nowadays, approximately 40% of the MRI exams are carried out after administration of a CA.The materials used as CAs for clinical studies must meet the following requirements: (1) high relaxivity; (2) high water solubility or colloidal stability; (3) low osmolality to avoid pain from osmotic shock and other adverse effects upon injection; (4) low toxicity, which requires both high thermodynamic and kinetic stability (no transmetallation by endogenous metal ions such as Zn 2+ ) in the case of metal chelates, and no leaching of toxic free metal ions in vivo from NPs; (5) rapid excretion, ideally not much longer than the MRI exam time; (6) a biodistribution with high specificity for the area of interest.
Various MRI contrast agents have been investigated and developed to achieve this goal.Gd-based CAs (GBCAs) include macrocyclic chelates of open-chain DTPA-type (DTPA = diethylenetriaminepentaacetic acid) or macrocyclic DOTA-type (DOTA = 2,2 ′ ,2 ′′ ,2 ′′′ -(1,4,7,10tetraazacyclododecane-1,4,7,10-tetrayl) tetraacetic acid) derivatives, such as Magnevist ® (Bayer Schering Pharma AG, Berlin, Germany), Dotarem ® (Guerbet, Paris, France), Omniscan ® (GE Healthcare, Chicago, IL, USA) and ProHance ® (Bracco, Milan, Italy) [7], which are clinically used as T 1 MRI CAs.These molecular agents have low r 1 and r 2 values and short blood circulation times due to their efficient renal excretion, and thus large amounts of injection doses are needed (typically 0.1 mmol/kg body weight) to achieve detectable contrast levels.This can increase the risk of toxicity due to potential release of free Gd 3+ ions in the body [14].This has occurred in some cases to patients with chronic kidney disease, which developed nephrogenic systemic fibrosis (NSF), characterized by skin thickening and hyperpigmentation and extracutaneous fibrosis [15][16][17].In addition, recent studies indicated that clinically developed GBCAs could be deposited in the brain after repeated use, and could cause neurotoxicity, although this has not been proven [18,19].In both cases, tissue deposits of linear GBCAs are much higher than those of macrocyclic GBCAs.
Magnetic NP (MNP)-based MRI CAs have many advantages relative to metal chelates, including larger magnetic moments and longer blood circulation times, which cause higher image contrast.They can also be used as multifunctional nanoplatforms for multimodal imaging, therapy and drug delivery (theranostics) after surface functionalization.[21].MNPs are composed of two parts: a magnetic core that enhances MRI and a surface-coating ligand layer responsible for colloidal stability and minimizes toxicity.This review focusses on recent developments and optimization and in vivo applications of MNPs containing Gd, Mn, Fe and other lanthanide ions which may function as dual-mode T 1 -T 2 MRI contrast agents (DMCAs).As the kind of contrast provided by MNPs depends fundamentally on their r 2 /r 1 ratios, the description of the principles for rational design of DMCAs starts by a brief summary of the basic theory of paramagnetic relaxation in MNPs.Then, as their r 2 /r 1 ratios are critically dependent on the MNPs composition, size and surfacecoating, the different ways in which these properties can be modulated to design T 1 or T 2 single-mode MRI CAs or T 1 -T 2 DMCAs are illustrated using selected examples.As the MNPs composition is a very important factor, a particular attention is given to systems using a single type of contrast material (e.g., Gd-or Mn-based T 1 contrast materials or superparamagnetic T 2 contrast materials) or both types of contrast materials in the same nanoplatform.

Basic Principles of Paramagnetic Relaxation in Small Complexes and Nanoparticles
The observed longitudinal proton relaxation rate (R obs 1 ) of a small paramagnetic complex in aqueous solution is the sum of a paramagnetic and a diamagnetic term, where the first is proportional to the concentration of the paramagnetic ion and the second is the contribution of the water solvent: The observed relaxivity (in mM −1 s −1 units) is the sum of the inner (r 1is ) and outer sphere (r 1os ) terms: The same equations apply to R 2 and r 2 .These contributions are evaluated using the dominant dipole-dipole relaxation mechanism, which is modeled by the Solomon-Bloembergen-Morgan (SBM) theory for r 1is , and the Freed theory for r 1os [3][4][5][6]22].For r 1 , the main parameters determining the IS contribution are the number of water molecules in the first coordination sphere of the metal ion (q), and the three processes responsible for the time fluctuation of the nucleus-electron interactions, which are the inner-sphere water residence time (τ M = k ex −1 , where k ex is the water exchange rate), the molecular reorientational correlation time (τ R ) and the electron spin relaxation time (T 1e ) of the metal ion, which is frequency dependent.The OS contribution is determined by the diffusion correlation time (τ D ) and the distance of closest approach (a) of the water molecules freely diffusing near the complex.
The r 1 and r 2 values of paramagnetic NPs also have in principle IS and OS contributions.The IS contribution results from the exchange of the water protons directly coordinated to the metal ions at the NPs surface with bulk water, as those located below the surface have a negligible effect.The IS r 1 depends on the hydration number of the surface ions and the NP's surface-to-volume ratio [23].In the case of Gd 3+ ions, they affect the bound water proton relaxation mainly through the dipolar mechanism, which can be modeled by the SBM equations.When other paramagnetic Ln 3+ ions with high magnetic moments and very short T 1e values (e.g., Tb 3+ , Dy 3+ , Ho 3+ ) (Table 2) are present, a Curie term is also present [24][25][26].* The magnetic moment of Eu 3+ is not zero due to the contribution of the excited states 5 D 0 and 5 D 1 .
However, for surface-coated NPs, the OS contribution dominates the T 1 and T 2 relaxation, as the water protons are indirectly in contact with the paramagnetic metal ions in a NP due to the surface coating.It results from the diffusion of water molecules in the fluctuating magnetic field inhomogeneities created in their vicinity by the magnetized NPs, and does not contain the Curie contribution.The r 1 value is approximately proportional to the square of the spin magnetic moment (µ s 2 = S(S + 1)h 2 for transition metal ions and µ s 2 = 4S(S + 1) + L(L + 1)h 2 for Ln 3+ ions, where S is the spin quantum number and L is the orbital quantum number) multiplied by the number (N) of Ln 3+ ions in a NP which can interact with a water proton, as given by Equation (3): Transverse relaxation of water protons is induced by fluctuations of local magnetic fields generated by the NPs.Thus, r 2 is proportional to the square of the total magnetic moment (µ) of the NP, as given by Equation (4): In the case of superparamagnetic NPs, the OS contribution is also dominant.However, for a colloidal dispersion in the presence of a magnetic field, the return of their magnetization to equilibrium is determined by two different processes [11]: (a) the Néel relaxation, describing the return of the magnetization of each of the NPs to equilibrium after a perturbation that tilts that magnetization away from the direction of its easy axis; its relaxation time (τ N ) defines the fluctuations that arise from jumps of the magnetization between different easy directions; (b) Brownian relaxation, defined by the relaxation time τ B , which characterizes the viscous rotation of the particle.The global magnetic relaxation rate of the colloid is the sum of the Néel (τ N −1 ) and Brownian (τ B −1 ) relaxation rates, , where τ is the global magnetic relaxation time.In these systems, r 1 and r 2 can be described by Freed's model for paramagnetic systems, using the diffusion correlation time (τ D ) and considering the electron spin longitudinal relaxation time τ S1 as equal to the Néel relaxation time τ N .
The magnetization of the superparamagnetic iron oxide NPs reaches its saturation value M S at B 0 ≤ 0.5 T, which is at the lower limit for commonly used clinical MRI scanners.Therefore, their R 2 is in practice usually independent of B 0 .The r 1 value is governed by the volume fraction of the superparamagnetic particles (υ), the diffusion correlation time (τ D = d 2 /4D, where d is the diameter of the particle and D is the diffusion coefficient), and the magnetization of the NP (M S ) at the B 0 value of the clinical MRI scanner.A discussion of the transverse relaxivity of spherical superparamagnetic NPs can be found in the literature [11,[27][28][29].
Superparamagnetic NPs are the T 2 CAs most commonly used in MRI, usually with a very high r 2 (typically 60-400 s −1 mM −1 ), in particular the SPIONs.This can hamper the interpretation of T 2w images due to the difficulty in distinguishing the CA-induced darkening from partial-volume artifacts, motion artifacts, and tissue inhomogeneities [30].

NPs for T 1 or T 2 Single-Mode MRI Contrast
The type of contrast provided by MNPs as MRI CAs depends on their r 1 and r 2 values, as well as their r 2 /r 1 ratios, which are dependent on their composition, size and surfacecoating [31].The ideal T 1 MRI CA should have high r 1 values and r 2 /r 1 ratios close to 1.0.The most important contribution to the r 1 values comes from the metal ions present on the MNP surface, which interact directly with nearby water proton spins by an IS mechanism.The most efficient paramagnetic ions are Gd 3+ (S = 7/2), Mn 2+ (S = 5/2), and Fe 3+ (S = 5/2) (S is the total spin quantum number) due to their high number of unpaired electrons, high magnetic moments and long T 1e values (Table 2) [3][4][5][6]22].
The use of high spin Fe 3+ ions as T 1 MRI probes within nanostructures for in vitro studies or preclinical investigations has been recently reviewed [52], including examples such as amphiphilic polymer-based NPs like Fe 3+ -chelated poly(lactic-co-glycolic) acid (PLGA) NPs, NPs containing polyphenolic Fe 3+ -binding units such as Fe 3+ -loaded synthetic melanin nanoparticles (SMNPs) and bimetallic (Gd, Fe)-phenolate coordination polymer (CN) NPs.In several of these systems, although the Fe 3+ ion is not hydrated (q = 0) and therefore the IS (q ≥ 1) contribution to relaxivity is not present, the reported r 1 values are higher than those expected for the OS (q = 0) contribution alone, which suggests the presence of a significant contribution from the water molecules of the second hydration sphere (SS).
The ideal T 2 MRI CAs should have high r 2 values and high (≥10) r 2 /r 1 ratios to induce predominantly T 2 proton spin relaxation, leading to a decrease in signal intensity in T 2w or T 2w * MRI images [31].The systems that fulfil these conditions are SPIONs and paramagnetic lanthanide (Ln 3+ )-based NPs (Ln = Tb, Dy and Ho), specially at high magnetic fields (B 0 ).
Paramagnetic NPs containing Tb 3+ , Dy 3+ or Ho 3+ have high r 2 values due to their high magnetic moments and very short (<1 ps) electronic relaxation times (Table 2) [24].Therefore, they can be alternatives to SPIONs as efficient T 2 and T 2 * CAs [13,[57][58][59][60], in particular at high MRI fields (B 0 > 3 T).This is because r 2 generally increases with B 0 , but the contribution from the Curie spin relaxation mechanism increases with B 0 2 .The maximum r 2 values are obtained at slower inner sphere water exchange (τ M = 0.1-10 µs) than for r 1 (τ M = 1-100 ns) [61].Also, in contrast to iron oxide particles, Ln 3+ -containing NPs show no saturation of the magnetization, even at magnetic field strengths as high as 30 T [62].
The highest payload of Ln 3+ ions per particle at a particular site can be delivered by inorganic Ln 3+ -containing NPs, including Ln 2 O 3 , LnF 3 , and NaLnF 4 .NPs with a diameter of 50-100 nm contain approximately 106 Ln 3+ ions each and their magnetic properties make them good candidates for T 2w and T 2w * MRI CAs [63].Dy 3+ -based nanomaterials are the most studied, as they show the largest T 2 relaxation effects [23].Moreover, the T 2 effects for Dy 3+ -based nanomaterials, such as β-NaDyF 4 , are two orders of magnitude higher than for clinically approved iron oxide nanomaterials under the increasingly used very high field magnets (7 T and higher).This arises due to the Curie spin relaxation mechanism of the Dy 3+ -based systems [64].However, coating of the particles is necessary to avoid leaching of free Ln 3+ .

NPs as T 1 -T 2 DMCAs
While conventional MRI CAs respond only in a single imaging mode, either T 1 or T 2 , NPs with multimodal capabilities provide complementary diagnostic information.However, when hybrid imaging systems, such as PET/CT or PET/MRI scanners are not available, a combination of two different imaging devices must be used separately, which is an inconvenient, time-consuming and expensive procedure [65][66][67][68][69].The development of MRI T 1 -T 2 dual-mode CAs in a single nanoplatform is an attractive solution to overcome ambiguities in conventional MRI diagnostics, especially when the biological targets are small, as well as the image matching difficulties caused by relocating the imaging object, and by the discrepancies resulting from different depth penetrations and spatial/time resolutions of multiple imaging strategies [69].In fact, T 1 -T 2 dual-modal MRI images can be easily acquired by changing the parameters of the pulse sequences in the operational mode of the same MRI scanner.
Dual-modal T 1 -T 2 MRI CAs (DMCAs) should have both high r 1 and r 2 values, with r 2 /r 1 ratios (~2-10) between those of ideal T 1 and T 2 MRI CAs.If Gd-based NPs are to be used as DMCAs, their r 2 /r 1 ratios should be increased from their usual values of ~1, while the use of SPION-based DMCAs requires a decrease in their r 2 /r 1 ratios from their usual very high values.Mn-based NPs are useful as DMCAs because of their suitable r 2 /r 1 ratios.
The approaches proposed to design DMCAs comprise: (1) use of NPs with a single type of contrast material, Gd-or Mn-based NPs or superparamagnetic NPs with appropriate size and magnetization to provide T 2 and T 1 contrast; or (2) include both T 1 and T 2 contrast materials in the same NP.Both have advantages and disadvantages [31,[69][70][71][72]
Finally, a Fe-based dual-mode T 1 -T 2 MRI DMCA was reported consisting of a onepot synthesized nanostructured coordination polymer containing Fe 3+ (Fe-NCP) functionalized with BSA.These NPs, with a hydrodynamic diameter of 97 nm and a ζ potential of −31.2 mV, had high colloidal stability, low cytotoxicity and good relaxivities, r 1 = 5.3 mM −1 .s−1 and r 2 = 10.9 mM −1 .s−1 (r 2 /r 1 = 2.1) at 7.0 T. Their T 1 -T 2 MRI contrast was verified in in vitro phantoms, ex vivo C57BL/6J mice and in vivo GL261 glioblastoma tumor-bearing mice.The vivo MRI of Fe−NCPs showed high T 1 and T 2 contrast in the tumor in a very short period of time and were safe for the mouse.Their large long-term uptake in the spleen was related to their rapid clearance by the mononuclear phagocytic system (MPS) due to the NPs size being bigger than 40 nm [83].
Table 3 shows that not all the systems described reach the low r 2 /r 1 values recommended for efficient DMCAs, illustrated by the second and fourth examples discussed above [77,83].

2.
NPs based on a typical T 2 agent: USPIONs and FeO x This strategy is based on using superparamagnetic NPs with appropriate size and magnetization to provide T 2 and T 1 contrast simultaneously.Although SPIONs with a core diameter less than 10 nm can produce positive contrast in T 1w images at low concentrations, their high T 2 effects (high r 2 /r 1 ) resulting from their high magnetic moment limit their application as T 1 -T 2 DMCAs.However, the magnetic moment of Fe 3 O 4 NPs is strongly dependent on their size and decreases rapidly as their size decreases due to the reduction in the volume magnetic anisotropy and spin disorder (canting) at their surface, which suppresses the T 2 effect and therefore maximizes the T 1 contrast effect, controlling their relaxivities [11].Therefore, the appropriate ultra-small-sized Fe 3 O 4 NPs (USPIONs) are potential candidates for T 1 -T 2 DMCAs.
Some examples of such systems are summarized in Table 4, where the USPION core was coated with hydrophilic polymers, e.g., poly(methacrylic acid) (PMAA)-polytrimethylene terephthalate (PMAA-PTTM) [85], poly(acrylic acid) (PAA) [86,87] and PEG [88], or silica [89], to prevent the aggregation of the NPs and to ensure a small particle size.The core was usually spherical, but in some cases nanoplates [89] or nanocubes [90,91] were used.Some of the systems summarized in Table 4 are now discussed in more detail.
Li et al. reported monodispersed water-soluble and biocompatible USPIONs (3.3 nm average diameter) grafted with thiol functionalized (PMAA)-(PTTM) using a high-temperature co precipitation method.These NPs, with r 1 = 8.3 and r 2 = 35.1 s −1 .mM−1 (r 2 /r 1 = 4.2) at 4.7 T, showed in vitro and in vivo potential as T 1 -T 2 DMCAs upon i.v.injection in mice, as positive and negative contrasts were observed in the T 1w and T 2w MRI images of liver and kidneys, respectively [85].Miao et al. developed PAA-coated USPIONs (5.1 nm core diameter and 41.35 nm average hydrodynamic size) with r 1 = 10.52 and r 2 = 38.97s −1 .mM−1 (r 2 /r 1 = 3.70) at 1.41 T. The NPs had in vitro and in vivo potential as T 1 -T 2 DMCAs, as illustrated by T 1w and T 2w MRI images at 3 T, with positive contrast observed in the rabbit vasculature, while the rabbit popliteal lymph node exhibited negative contrast [87].Wang et al. reported USPION-PEG (P-UDIOC) NPs with extremely small core size (2.3 nm) and a compact hydrophilic PEG surface and with r 1 = 1.37 and r 2 = 7.53 s −1 .mM−1 (r 2 /r 1 = 5.5) at the ultrahigh field (UHF) of 7.0 T. These UHF-tailored T 1 -T 2 DMCAs showed dual enhanced T 1 -T 2 contrast at 7 T and enabled a clear visualization of microvasculature as small as ≈140 µm in diameter under UHF MRI (Figure 3), extending the detection limit of 7 T MRI angiography (MRA) [88].
An example of a system with a non-spherical core consisting of Fe 3 O 4 crystal nanoplates was reported by Zhou et.al. [89] These superparamagnetic magnetite nanoplates had (111) exposed facets which much contributed to their T 1 and T 2 contrast effects.The main contribution to r 1 of the magnetic nanoplates is the chemical exchange with bulk water (given by the exchange lifetime, τ M ) of the water molecules bound at the inner-sphere of the Fe 3+ ions on the highly exposed particle surface iron-rich Fe 3 O 4 (111) surfaces, according to the dominant IS model [11].The r 2 values are dominated by the OS mechanism, which accounts for the effect of fluctuations of the local magnetic field inhomogeneities induced by the tumbling NPs on the protons of outer-sphere water molecules diffusing nearby.According to Freed's theory, the r 2 value is proportional to the square of the magnetization (M s 2 ), which relates to the intrinsic superparamagnetism of the nanoplates and determines the strength of the local magnetic field inhomogeneities, as well as to the square of the effective radius of the magnetic core (R 2 ), which determines the field perturbation areas for the outer-sphere protons.The rapid random flipping of the anisotropic nanoplates, when represented by an equivalent simulated sphere, generate a larger area of local field inhomogeneity compared with nanospheres under the same applied magnetic field.The balance of T 1 and T 2 contrasts was attained by controlling the structure and surface features of the nanoplates, including morphology (nanoplates vs. spheres in IOP-4.8@stPE) and surface coating (e.g., IOP-4.8 vs. IOP-4.8@SiO 2 and IOP-4.8@stPE), with a large decrease in r 1 due to blocking of the exposure of the facets causing the disappearance of the IS contribution and corresponding decrease in r 2 /r 1 , leading to change from DMCAs to T 2 agents.The decrease in the nanoplate thickness decreased the r 2 /r 1 value, as IOP-8.8(r 2 /r 1 ~8.18) is T 2 -dominated, while the IOP-4.8(r 2 /r 1 ~4.22) is a T 1 -T 2 DMCA [89].An example of a system with a non-spherical core consisting of Fe3O4 crystal nanoplates was reported by Zhou et.al. [89] These superparamagnetic magnetite nanoplates had (111) exposed facets which much contributed to their T1 and T2 contrast effects.The main contribution to r1 of the magnetic nanoplates is the chemical exchange with bulk water (given by the exchange lifetime, τM) of the water molecules bound at the innersphere of the Fe 3+ ions on the highly exposed particle surface iron-rich Fe3O4 (111) surfaces, according to the dominant IS model [11].The r2 values are dominated by the OS mechanism, which accounts for the effect of fluctuations of the local magnetic field inhomogeneities induced by the tumbling NPs on the protons of outer-sphere water molecules diffusing nearby.According to Freed's theory, the r2 value is proportional to the square of the magnetization (Ms 2 ), which relates to the intrinsic superparamagnetism of the nanoplates and determines the strength of the local magnetic field inhomogeneities, as well as to the square of the effective radius of the magnetic core (R 2 ), which determines the field perturbation areas for the outer-sphere protons.The rapid random flipping of the anisotropic nanoplates, when represented by an equivalent simulated sphere, generate a larger area of local field inhomogeneity compared with nanospheres under Table 4 shows that not all the systems described reach the low r 2 /r 1 values recommended for efficient DMCAs [90,91].

Triggered aggregation change in ESIONs
The design of T 1 -T 2 DMCAs can also be based on a different mechanism, based on the change in the aggregation state of extremely small-sized iron oxide nanoparticles (ESIONs) in the size range of 1.5-4 nm with relatively high r 1 values, triggered by an external signal, such as pH change, a redox reaction or laser light.The aggregation of ESIONPs can generate T 2 contrast effects due to the enhancement of magnetic field inhomogeneity and magnetic coupling between Fe centers.Some examples of such systems are summarized in Table 4.An example is based on pH-sensitive hydrazine functionalized ESIONPs forming assemblies (IONAs) cross-linked by small-molecular aldehyde derivative ligands.The dynamic formation and cleavage of hydrazone (-C=N-N-) (f) linkages in neutral and acidic environments, respectively, is the base for the reversible response of the nanoassemblies to pH variations.At pH 7.5, IONAs are stable and with high r 2 /r 1 = 34.2,while at acidic pH 5.5, such as in the acidic tumor microenvironment (TME), the hydrazone bonds are cleaved and the IONAs are disassembled into hydrophilic dispersed ESIONs, with a r 2 /r 1 = 4.1.The change in this relaxivity ration results from an increase in the number of second sphere water molecules (q SS ) and a large decrease in the magnetization (M s ) of the dispersed ESIONs relative to the IONAs, which affect their r 1 and r 2 values, respectively [92].Another example is based on ESIONs linked with the targeting ligand folic acid (FA) binding arthritis-associated macrophage cells, and the light responsive diazirine (DA) through a PEG spacer (Fe 3 O 4 -PEG-(DA)-FA).These nanoparticles can form nanocomposites (NCs) upon laser irradiation to have tunable r 1 and r 2 values upon variation in the laser irradiation time.The change in the r 2 /r 1 value from 2.36 in the NPs to 18.8 in the NCs led to the use of the designed Fe 3 O 4 -PEG-(DA)-FA NPs as T 1 CAs in in vivo MRI of an arthritis mouse model without lasers and enhanced T 1 -T 2 DMCAs in the arthritis inflamed region under laser irradiation due to the formation of NCs that accumulated within the arthritis region and their limited intravasation back to the blood circulation [93].
Again, Table 4 shows that not all the systems described reach the low r 2 /r 1 values recommended for efficient DMCAs: while the first one does [92], the second one reaches quite high r 2 /r 1 values [93].Besides this, in some cases, T 1 /T 2 triggering effects, such as pH decrease or reduction by GSH in the TME, led to final r 2 /r 1 values below or above the range where T 1 -T 2 DMCAs operate, forming instead T 1 /T 2 switching MRI CAs [94][95][96][97][98][99][100][101].This illustrates how the control of the switching mechanism at the molecular level is difficult.* Not available.

DMCAs including Both T 1 and T 2 Contrast Materials in the Same Nanoparticle
The strategy of including both T 1 and T 2 contrast materials in the same nanoplatform has used several different designs to obtain hybrid nanostructures: (a) NPs integrating a superparamagnetic T 2 -contrasting material inside a paramagnetic T 1 material (e.g., a Gd 3+ complex); (b) doping a superparamagnetic NP with paramagnetic T 1 contrast mate-rials inside them; (c) T 1 and T 2 contrast materials connected side-by-side to form hybrid nano-oligomers (DB-HNT).In all these designs, with both contrast materials in the same nanoplatform, the magnetic fields generated by each CA disturb the relaxation process of the other.However, as typical paramagnetic T 1 CAs have low r 1 values compared to the r 2 values of superparamagnetic T 2 agents due to their relative magnetic moments, their effect on the T 1 relaxation of the T 1 contrast material is much larger than the opposite.The T 2 agent generates a strong magnetic field induced by the external B o , which is dependent on 1/r 3 (r = distance from the T 2 agent), which affects the electronic spin relaxation time (T 1e ) of the paramagnetic T 1 agent depending on their relative locations.The resulting effects of this magnetic coupling depends on the separation distance (d) between T 1 and T 2 agents [102] and also on their relative positions within the NP (Figure 4) [103].
* Not available.

DMCAs Including Both T1 and T2 Contrast Materials in the Same Nanoparticle
The strategy of including both T1 and T2 contrast materials in the same nanoplatform has used several different designs to obtain hybrid nanostructures: (a) NPs integrating a superparamagnetic T2-contrasting material inside a paramagnetic T1 material (e.g., a Gd 3+ complex); (b) doping a superparamagnetic NP with paramagnetic T1 contrast materials inside them; (c) T1 and T2 contrast materials connected side-by-side to form hybrid nano-oligomers (DB-HNT).In all these designs, with both contrast materials in the same nanoplatform, the magnetic fields generated by each CA disturb the relaxation process of the other.However, as typical paramagnetic T1 CAs have low r1 values compared to the r2 values of superparamagnetic T2 agents due to their relative magnetic moments, their effect on the T1 relaxation of the T1 contrast material is much larger than the opposite.The T2 agent generates a strong magnetic field induced by the external Bo, which is dependent on 1/r 3 (r = distance from the T2 agent), which affects the electronic spin relaxation time (T1e) of the paramagnetic T1 agent depending on their relative locations.The resulting effects of this magnetic coupling depends on the separation distance (d) between T1 and T2 agents [102] and also on their relative positions within the NP (Figure 4) [103].1. NPs with a T2 material inside a paramagnetic T1 material For systems consisting of a superparamagnetic core within a paramagnetic shell, the strong magnetic field from the core opposes the magnetic field created by the shell and reduces it (Figure 4, left), strongly quenching its r1 value.Their magnetic interaction is proportional to the inverse sixth power of their separation distance (d −6 ), as shown by a study where the distance-dependent magnetic resonance tuning (MRET) strategy for tuning the r1 value of a T1 agents was introduced [102].A series of spherical NPs was designed, consisting of three components, where the paramagnetic enhancer (Gd-DOTA)

1.
NPs with a T 2 material inside a paramagnetic T 1 material For systems consisting of a superparamagnetic core within a paramagnetic shell, the strong magnetic field from the core opposes the magnetic field created by the shell and reduces it (Figure 4, left), strongly quenching its r 1 value.Their magnetic interaction is proportional to the inverse sixth power of their separation distance (d −6 ), as shown by a study where the distance-dependent magnetic resonance tuning (MRET) strategy for tuning the r 1 value of a T 1 agents was introduced [102].A series of spherical NPs was designed, consisting of three components, where the paramagnetic enhancer (Gd-DOTA) was separated from the superparamagnetic quencher (a 12 nm Zn 0.4 Fe 2 .6 O 4 NP) by controlling the thickness of a SiO 2 separating layer decreasing from 18 to 2 nm (Zn 0.4 Fe 2 .6 O 4 @SiO 2 @ Gd-DOTA).A decrease in r 1 was consistently observed as the separation distance between the T 1 and T 2 agents decreased.The T 1 MRI signal was quenched when the d value between the enhancer and the quencher decreased and r 1 decreased from 1.58 to 0.13 mM −1 .s−1 (3 T), which was a consequence of the increase in the T 1e value of the T 1 agent [102].
The strategy of including a layer of increasing thickness to increase the T 2 core-T 1 shell distance, and thus modulating their magnetic coupling, was pursued using not only inorganic porous materials like SiO 2 , but also micellar structures incorporating organic block copolymers and inorganic porous materials as possible frameworks [104].The basic properties of several examples from the literature are summarized in Table 5.These involve the formation of spherical or cubic core-shell NPs integrating a superparamagnetic T 2 -contrasting core (e.g., Fe 3 O 4 , MnFe 2 O 4 ) inside a paramagnetic T 1 shell (e.g., Gd 2 O 3 , Gd 2 O(CO 3 ) 2 , MnO) [105][106][107][108][109][110][111][112], or conjugating the superparamagnetic core with a paramagnetic Gd 3+ or Mn 2+ complex at its surface [113][114][115][116][117]. Generally, a sharp decrease in the magnetic coupling effect upon increasing the core-shell separation was verified experimentally by including a silica shell of increasing thickness between the two materials [105][106][107], or by increasing the distance between the core and the layer of pendant paramagnetic complexes through longer spacer groups [102,113].Here, we will discuss only a few of the examples in more detail.
For instance, a core-shell-type T 1 -T 2 DMCA agent has been described, where the T 1 contrast material, a Gd 2 O(CO 3 ) 2 layer (1.5 nm thickness) was located on the shell in order to be in direct contact with water molecules, to obtain for high T 1 contrast; the superparamagnetic T 2 contrast material, MnFe 2 O 4 (15 nm size) was located at the core, from where it could induce a long-range magnetic field for the relaxation of water molecules.The two materials were separated by a SiO 2 layer of increasing (4, 8, 12, 16 and 20 nm) thickness (MnFe 2 O 4 @SiO 2 @Gd 2 O(CO 3 ) 2 NPs).As the SiO 2 layer became thicker, the magnetic coupling decreased, the T 1 quenching was reduced and r 1 increased (from 2.0 to 32.5 mM −1 .s−1 ) while the r 2 decrease was weaker (332 to 213 mM −1 .s−1 ).When the thickness of the SiO 2 layer was equal of larger than 16 nm, the r 2 /r 1 values decreased from 160 to 6.5, and the NPs became T 1 -T 2 DMCAs, as both T 1 and T 2 effects became larger than the effects of the individual single-mode contrasts (Figure 5) [105].
ic block copolymers and inorganic porous materials as possible frameworks [104].The basic properties of several examples from the literature are summarized in Table 5.These involve the formation of spherical or cubic core-shell NPs integrating a superparamagnetic T2-contrasting core (e.g., Fe3O4, MnFe2O4) inside a paramagnetic T1 shell (e.g., Gd2O3, Gd2O(CO3)2, MnO) [105][106][107][108][109][110][111][112], or conjugating the superparamagnetic core with a paramagnetic Gd 3+ or Mn 2+ complex at its surface [113][114][115][116][117]. Generally, a sharp decrease in the magnetic coupling effect upon increasing the core-shell separation was verified experimentally by including a silica shell of increasing thickness between the two materials [105][106][107], or by increasing the distance between the core and the layer of pendant paramagnetic complexes through longer spacer groups [102,113].Here, we will discuss only a few of the examples in more detail.
For instance, a core-shell-type T1-T2 DMCA agent has been described, where the T1 contrast material, a Gd2O(CO3)2 layer (1.5 nm thickness) was located on the shell in order to be in direct contact with water molecules, to obtain for high T1 contrast; the superparamagnetic T2 contrast material, MnFe2O4 (15 nm size) was located at the core, from where it could induce a long-range magnetic field for the relaxation of water molecules.The two materials were separated by a SiO2 layer of increasing (4, 8, 12, 16 and 20 nm) thickness (MnFe2O4@SiO2@Gd2O(CO3)2 NPs).As the SiO2 layer became thicker, the magnetic coupling decreased, the T1 quenching was reduced and r1 increased (from 2.0 to 32.5 mM −1 .s−1 ) while the r2 decrease was weaker (332 to 213 mM −1 .s−1 ).When the thickness of the SiO2 layer was equal of larger than 16 nm, the r2/r1 values decreased from 160 to 6.5, and the NPs became T1-T2 DMCAs, as both T1 and T2 effects became larger than the effects of the individual single-mode contrasts (Figure 5) [105].Schematic image of the core-shell-type dual-mode nanoparticle [MnFe2O4@SiO2@Gd2(CO3)2].The T1 contrast material is positioned on the shell to have direct contact with the water for high T1 contrast effects, and the superparamagnetic contrast material is located at the core, inducing a long-range magnetic field for the relaxation of water.Reproduced from [31].
A smart nanotheranostic system for early diagnosis and therapy of cancer was developed, consisting of camptothecin (CPT)-loaded mesoporous silica nanoparticles (MSN) capped with manganese oxide (MnOx)-coated SPIONs (MnOx- Schematic image of the core-shell-type dual-mode nanoparticle [MnFe 2 O 4 @SiO 2 @Gd 2 (CO 3 ) 2 ].The T 1 contrast material is positioned on the shell to have direct contact with the water for high T 1 contrast effects, and the superparamagnetic contrast material is located at the core, inducing a long-range magnetic field for the relaxation of water.Reproduced from [31].
A smart nanotheranostic system for early diagnosis and therapy of cancer was developed, consisting of camptothecin (CPT)-loaded mesoporous silica nanoparticles (MSN) capped with manganese oxide (MnO x )-coated SPIONs (MnO x -SPION@MSN@CPT NPs).The acid, oxidative stress and redox (GSH) response of MnO x regulated the CPT drug release from the MSN channels, while the high magnetization of the surface SPIONs achieved high r 2 values (102.2 mM −1 .s−1 ).At the same time, degradation of the MnO x shell caused release of Mn 2+ in the TME, enhancing r 1 (2 → 13.6 mM −1 .s−1 ).The efficacy of this MRI responsive theranostic T 1 -T 2 DMCA was confirmed in vitro on pancreatic cancer cells and in vivo on tumor-bearing mice (Figure 6) [111].
Table 5 shows that many of the systems described using this strategy have r 2 /r 1 values above the values recommended for efficient DMCAs [107,109,110,113,115,118,120,121].Besides this, the systems based on the release of toxic free Mn 2+ in the TME, although interesting for animal studies, are not suited for clinical applications [110,111].* Not available.

Doping superparamagnetic T 2 contrast NPs with paramagnetic T 1 contrast materials inside
For systems where the paramagnetic material resides inside the superparamagnetic iron oxide, the magnetic fields of both materials reinforce each other simultaneously, strongly enhancing the r 1 value and causing a synergistic T 1 -T 2 enhancement effect (Figure 4, right) [103].The basic properties of systems using this strategy are shown in Table 6, some of which are discussed in some detail.The theory of synergistic T 1 -T 2 enhancement effect discussed above was confirmed by Gao et al. using Gd 2 O 3 -embedded Fe 3 O 4 -HDA-G 2 NPs (GdIO-HDA-G 2 ), which showed a synergistic enhancement of r 1 and r 2 .The GdIO had higher r 2 (146.5 mM −1 .s−1 ) than Fe 3 O 4 (125.4mM −1 .s−1 ) of similar size and also higher r 1 (69.5 mM −1 .s−1 ) than Gd 2 O 3 (12.1 mM −1 .s−1 ) of similar size.Furthermore, the Gd 2 O 3 NPs showed no enhanced T 2 contrast, while Fe 3 O 4 nanoparticles showed limited enhanced T 1 contrast.Simultaneous in vivo T 1w and T 2w MRI of BALB/c mice upon i.v.injection of GdIO showed simultaneous strong MRI contrast enhancement of liver in both types of images due to the high accumulation of NPs in the hepatic Kupffer cells of the liver mononuclear phagocyte system (MPS).The same MRI experiment in HepG2 tumor mice detected the liver lesions through pseudo-negative and pseudo-positive contrast effects because the contrast between lesions and surrounding normal liver tissue increased due to the very low uptake by hepatic tumors, which contain few active Kupffer cells and macrophages (Figure 8).This work validated the new strategy for the design of new T 1 -T 2 DMCAs [103].Table 6.Summary of basic properties of NPs made of a paramagnetic T 1 material inside a superparamagnetic T 2 material or forming hybrid oligomers of different shapes (in bold).Their properties are compared with those of the NPs made of their components.Another example of the use of the same design strategy consisted of water-dispersible GdIO NPs stearic acid modified low molecular weight polyethyleneimine (stPEI) (GdIO-stPEI).This nanoplatform was capable of binding and delivering siRNA for gene knockdown and work as T 1 -T 2 DMCA with HCT-116 cells in vitro [123].Zwitterion dopamine sulfonate-coated superparamagnetic GdIO NPs (GdIO-ZDS) with small core size (2.8-4.8 nm) showed partial paramagnetism at room temperature, with a decreased M S value relative to IOs of the same core size.This resulted from the combination of surface canting with the effect of the embedded Gd 2 O 3 nanoclusters which disturbs the long-range order of magnetic spins in the small GdIO NPs.It also led to a significantly increased r 1 value and decreased r 2 value relative to IO-ZDS of the same core size.For example, GdIO NPs with a 4.8 nm diameter, had a high r 1 = 7.85 mM −1 .s−1 and a low r 2 /r 1 = 5.24 relative to IO (r 2 /r 1 = 9.59).These NPs caused a strong positive tumor contrast effect in T 1w MRI images of SKOV3 human ovarian tumor mice through an enhanced permeation and retention (EPR) effect [124].
Molecules 2024, 29, x FOR PEER REVIEW 17 of 28 fects because the contrast between lesions and surrounding normal liver tissue increased due to the very low uptake by hepatic tumors, which contain few active Kupffer cells and macrophages (Figure 8).This work validated the new strategy for the design of new T1-T2 DMCAs [103].Superparamagnetic Fe 3 O 4 NPs have also been doped with other paramagnetic ions, such as Mn 2+ (MnIO) and Eu 3+ (EuIO) [125,126].The MRI contrast abilities of uniform Mn 2+ -doped iron oxide (MnIO) NPs nanoparticles with 5, 7, 9 and 12 nm size were studied.The NPs were superparamagnetic at 300 K, with M s values which decreased with the decrease in the MnIO NP sizes, from 71.0 emu g −1 for the 12 nm NPs to 39.7 emu g −1 for the 5 nm NPs due to the spin canting effect at their surface.Their r 1 and r 2 values were highly size-dependent, with r 2 /r 1 values decreasing from 7.4 for the 12 nm NPs to 2.6 for the 5 nm NPs.Thus, by controlling the size of the MnIO NPs, T 1 -dominated, T 2dominated, and T 1 -T 2 DMCAs could be obtained with much higher contrast enhancement than the corresponding conventional iron oxide nanoparticles, as verified by in vivo MRI of BALB/c mice [125].Finally, Eu 3+ -dopped iron oxide nanocubes (EuIO) were developed as T 1 -T 2 DMCAs for in vivo MRI.The EuIO nanocubes were composed of mixed Fe 3 O 4 (magnetite) and Eu 2 O 3 nanoclusters of 10.0, 14.0 and 20.1 nm size, due to the large ionic radius of Eu 3+ ions (94.7 pm) which prevented them from occupying either the tetrahedral or the octahedral interstitial sites in the spinel structure.The EuIO nanocubes are partially paramagnetic at 300 K, which is different from the superparamagnetism of magnetite NPs, due to increased spin canting on their surface layer after Eu 2 O 3 embedding.The Eu 2 O 3 clusters located inside the iron oxide NPs nanoparticles disturbed the local magnetic field intensity of the whole NPS and reduced their M s values (~39.6 emu g −1 ) relative to magnetite NPs with a similar size (~53.4emu g −1 ) at 300 K.The larger EuIO NPs had higher M s values due to the loss of the spin canting effect on the particle surface.As a result, both r 1 and r 2 values of EuIO nanocubes could be tuned by varying their sizes and Eu doping ratios.Larger EuIO nanocubes had higher r 1 and r 2 values.The Eu/Fe molar ratio also had an important role in the r 1 and r 2 values of EuIO nanocubes: raising the Eu molar ratio increased r 1 due to the spin order of Eu 3+ ions which had the same orientation as the local magnetic field, while it decreased r 2 values due to the reduction in the M s values after Eu embedding.For instance, EuIO nanocubes of 14 nm diameter showed a high r 1 = 36.8mM −1 .s−1 , which is approximately 3 times higher than that of Fe 3 O 4 NPs (12.47 mM −1 .s−1 ) of similar size.After citrate coating, EuIO nanocubes produced enhanced T 1 and T 2 MRI contrast effects in Sprague Dawley rats as models for in vivo MRI studies, in particular in the cardiac and liver regions [126].
In summary, Table 6 shows that the r 2 /r 1 values of the composite NPs (in bold) are more suitable as DMCAs when compared with those of the NPs made of their components.

3.
T 1 and T 2 contrast materials connected side-by-side forming hybrid oligomers of different shapes A third strategy to control the interference by magnetic coupling between T 1 and T 2 materials present in a single-composite nanostructure is to engineer the architecture of heterogeneous NPs forming hybrid trimers and oligomers of different geometries, like dumbbell-shaped NPs [109,127] or nanoflowers [109] (Table 6).Dumbbell-shaped NPs, or so-called 'Janus' NPs, were synthesized, with two different components within one single structure, as solid-state analogues of bifunctional organic molecules to construct hybrid nanotrimers (HNTs) (Figure 9a, right panel), in which iron oxide and Au nanocrystals were connected by a platinum nanocube (Au-IONP).The surface of its Au component was covalently immobilized with Gd-cystamine-DOTA 2 (dithiol derivative of DOTA) as a T 1 material (Gd(DOTA)-HNTs).To reduce its magnetic coupling with the IONP (T 2 material), the size of the Au nanocrystals was increased by controlling the seed-mediated growth processes during the synthesis and the size of the Pt cubes was also increased to increase the distance (D, Figure 9a, right panel) between the IONP and the Au nanocrystals.The resulting heterotrimers with large Au crystals had a dumbbell structure, (Gd(DOTA)-DB-HNTs and Gd(DOTA)-XDB-HNTs, the latter with a larger Au component and thus a larger number of Gd per single NP).The IO component was covered with PEG chains to make the whole NPs water-soluble and biocompatible (Figure 9b).The Gd(DOTA)-DB-HNTs and Gd(DOTA)-XDB-HNTs had increased r 1 values due to the reduced magnetic coupling between the T 1 and T 2 components as their D values increased.Their r 2 values were similar due to the similar sizes and shapes of the iron oxide components.The calculation of the r 1 relaxivity of each particle was based on different concentrations of Gd: r 1 (mM [Fe + Gd] −1 .s−1 ) or r 1 ′ (mM [Gd] −1 .s−1 ).Even though the r 2 /r 1 ratio of Gd(DOTA)-DB-HNTs was 33, their r 2 /r 1 ´was only 4.2, indicating that they could be T 1 -T 2 DMCAs.Although Gd(DOTA)-XDB-HNTs have slightly higher r 1 ′ (32.1 mM [Gd] −1 .s−1 ), they did not have a better r 2 /r 1 ´ratio (4.2) than Gd(DOTA)-DB-HNTs because of their larger size of the IO component, which increased r 2 .The hybrid heterotrimers were highly stable in physiological conditions and induced simultaneous positive and negative contrast enhancements in in vivo MRI images of HT-29 tumor-bearing mice upon i.v.injection of Gd(DOTA)-DB-HNTs, showing their potential as T 1 -T 2 DMCAs [127].could be T1-T2 DMCAs.Although Gd(DOTA)-XDB-HNTs have slightly higher r1′ (32.1 mM [Gd] −1 .s−1 ), they did not have a better r2/r1´ ratio (4.2) than Gd(DOTA)-DB-HNTs because of their larger size of the IO component, which increased r2.The hybrid heterotrimers were highly stable in physiological conditions and induced simultaneous positive and negative contrast enhancements in in vivo MRI images of HT-29 tumor-bearing mice upon i.v.injection of Gd(DOTA)-DB-HNTs, showing their potential as T1-T2 DMCAs [127].10).All the NPs had a high r 2 /r 1 ratio and the dumbbell shaped Fe 3 O 4 /MnO NPs produced a negative contrast effect in T 2 * w MRI images and a positive contrast effect in T 1w MRI images upon releasing Mn 2+ ions in a low pH environment, in vitro (aqueous phantoms) and in vivo (normal brain of a nude mouse) obtained after i.v.injection.The same study using an orthotopic xenograft model of human hepatocellular carcinoma (HCC) showed high contrast between relatively hyperintense HCC and hypointense background liver parenchyma in T 2w MRI HCC and hypointense background liver parenchyma due to the presence of Kupffer cells in the later, while in T 1w MRI images a bright signal of the HCC tumor was observed after injection, due to HCC cell uptake of the Mn 2+ ions liberated from the NPs in the low pH environment of the TME (Figure 11).This induced organ-specific contrast enhancement showed the potential of the hybrid NPs and T 1 -T 2 DMCAs [109].Fe3O4/MnO hybrid nanocrystals were prepared based on seed-mediated growth of MnO on the surface of a Fe3O4, where the resulting structures depended on the size of the seed NP, producing core/shell spherical Fe3O4 (5 nm)@MnO, dumbbell shaped Fe3O4 (11 nm)/MnO, and flower-shaped Fe3O4 (21 nm)/MnO hybrid NPs from 5 nm, 11 nm, and 21 nm Fe3O4 NPs, respectively (Figure 10).All the NPs had a high r2/r1 ratio and the dumbbell shaped Fe3O4/MnO NPs produced a negative contrast effect in T2 * w MRI images and a positive contrast effect in T1w MRI images upon releasing Mn 2+ ions in a low pH environment, in vitro (aqueous phantoms) and in vivo (normal brain of a nude mouse) obtained after i.v.injection.The same study using an orthotopic xenograft model of human hepatocellular carcinoma (HCC) showed high contrast between relatively hyperintense HCC and hypointense background liver parenchyma in T2w MRI HCC and hypointense background liver parenchyma due to the presence of Kupffer cells in the later, while in T1w MRI images a bright signal of the HCC tumor was observed after injection, due to HCC cell uptake of the Mn 2+ ions liberated from the NPs in the low pH environment of the TME (Figure 11).This induced organ-specific contrast enhancement showed the potential of the hybrid NPs and T1-T2 DMCAs [109].In summary, the synthetically challenging strategy of designing DMCAs based on forming hybrid oligomers of different shapes using T 1 and T 2 contrast materials, although scientifically original, has so far produced nanosystems with too high r 2 /r 1 values [109,127].

General Issues of In Vivo Use of NPs as MRI Contrast Agents
Besides the problems of design and contrast efficacy described in previous sections, the use of MNPs as MRI CAs in vivo involves many other issues, including their colloidal stability, biocompatibility, absorption, distribution, metabolism, and excretion (ADME) [128], toxicity, targeting and potential theranostic use.
Several intrinsic properties of MNPs, such as size, shape, coating, charge and presence of surface ligands, influence their biodistribution, elimination and target site accumulation.MNPs should be colloidally stable in aqueous media and in body fluids, as in vivo agglomeration and precipitation causes unacceptable safety issues and limits their highperformance MRI function.The colloidal stability is ensured by high zeta potentials and coating with hydrophilic and biocompatible ligands, such as PEG, which also maintains their non-toxicity.The biocompatibility of the materials used in DMCAs is also quite variable.Fe 3+ -based NPs are more biocompatible than those containing Gd 3+ , Dy 3+ , Ho 3+ , Tb 3+ and Mn 2+ , because iron is an essential element in the human body, participating in a wide variety of metabolic processes, including oxygen transport, DNA synthesis, and electron transport.However, its concentration in body tissues must be tightly regulated because in excessive amounts, Fe 2+ is toxic through the generation of reactive oxygen species (ROS) via the so-called Fenton reaction, which eventually results in cell damage, cell death, and organ failure, primarily affecting the liver, heart, pancreas, thyroid, and central nervous system [129].Although Mn is also an essential trace element, excessive doses of Mn 2+ can be neurotoxic, and its build-up in the brain may lead to manganism, a neurological disorder similar to Parkinsonism [130].Also, free Gd 3+ ions liberated from Gd chelates or NPs into the body could cause nephrogenic systemic fibrosis (NSF) [14,15].Non-toxicity of metal-based NPs is critical for their in vivo use as MRI CAs [131].
Because MRI CAs are usually intravenously injected, the preferable excretion of MNPs is through the renal system rather than the hepatobiliary pathway because the later one is relatively slow and the MNPs could be taken up by the liver reticuloendothelial system (RES) and be partially decomposed during the excretion process, which would be toxic to the body.This metabolization process of small (<10 nm) NPs can occur through direct penetration of the cell membrane (phagocytosis) and delivery to the cytoplasm and nucleus, while for larger NPs (10-200 nm) the uptake is mediated by clathrin-dependent endocytosis-Both mechanisms facilitate the cell lysosomal internalization and degradation of the MNPs [128].
For renal excretion, MNPs should be ultrasmall with hydrodynamic diameters less than 5 nm because the glomerular filtration diameter in the kidneys is 4.5-5 nm [132].The kinetic stability of MNPs should also be high to avoid their decomposition until they are excreted as urine through the renal system.
In summary, MNPs as MRI CAs for safe in vivo applications should be kinetically stable (i.e., no decomposition), coated with hydrophilic and biocompatible polymers for non-toxicity and colloidal stability, and ultrasmall with hydrodynamic diameters less than 5 nm for renal excretion.This excretion route, avoiding as far as possible their liver retention in the hepatobiliary pathway, is also very important for targeting and theranostic drug delivery by MNPs [133,134].In this review, several examples of in vivo use of theranostic DMCAs were presented, e.g., using ferroptosis inhibition [78], immunotherapy [79] and photodynamic therapy (PTT) associated with chemotherapy (ChT) [82] and chemodynamic therapy (CDT) [84].As far as nanocarriers for drug delivery and theranostics are concerned, the inappropriate release, internal instability and tissue non-targeting effects, sometimes observed due to biological barriers, are the main restrictions for their in vivo application [128,135].

Conclusions
This review describes recent developments and optimization of MNPs containing Gd, Mn, Fe and other lanthanide ions as potential dual-mode T 1 -T 2 MRI contrast agents (DMCAs).Their high performance was highlighted by describing selected in vivo MRI studies.However, the development stage of most of the reported MNP-based MRI CAs is still quite limited, both in vitro and in preclinical in vivo small animal studies.To improve their chances to reach the clinical trials stage, several key issues must be solved.These include long-term colloidal stability in aqueous media, toxicity and pharmacokinetics, which depend on their coating with hydrophilic and biocompatible ligands.In comparison with MNPs containing Mn, or lanthanides (such as Gd, Dy, Ho, and Tb), Fe-based NPs are less toxic because iron is an essential element, and several IONPs have been approved and commercialized as MRI CAs, such as Feridex ® , Sinerem ® , and Resovist ® , before being discontinued due to lack of commercial interest.The understanding of the correlation between the physicochemical properties of MNPs and their biological behavior in vivo must also be improved using appropriate preclinical small animal studies.
Another major problem faced by the clinical introduction of new MRI CAs in general, and MNP-based MRI CAs in particular, due to their molecular complexity and the consequent expensive synthesis, is to solve the main regulatory challenges of controlled NPs synthesis, uniformity, batch-to-batch reproducibility and upscaling of production.This leads to large development costs, especially for phase III studies, compared with the post-approval revenues [136,137].However, the large amount of information obtainable with molecular imaging using MNP-based MRI CAs, such as those described in this review, may contribute to early diagnosis and personalized therapies, which ultimately will reduce health care costs. .

Figure 2 .
Figure 2. (Left): a model structure of Mn-DMSS NPs; (Right): Simultaneous T1 and T2 imaging of a SD rat liver on a 3.0 T MRI scanner: In vivo T1-weighted (top) and T2-weighted (bottom) MR images.Reproduced with permission from [77].Copyright from the American Chemical Society.

Figure 2 .
Figure 2. (Left): a model structure of Mn-DMSS NPs; (Right): Simultaneous T1 and T2 imaging of a SD rat liver on a 3.0 T MRI scanner: In vivo T1-weighted (top) and T2-weighted (bottom) MR images.Reproduced with permission from [77].Copyright from the American Chemical Society.

Figure 2 .
Figure 2. (Left): a model structure of Mn-DMSS NPs; (Right): Simultaneous T 1 and T 2 imaging of a SD rat liver on a 3.0 T MRI scanner: In vivo T 1 -weighted (top) and T 2 -weighted (bottom) MR images.Reproduced with permission from[77].Copyright from the American Chemical Society.

28 Figure 3 .
Figure 3. Vascular imaging performance of P-UDIOC-enhanced UHF MRI.(a) P-UDIOC, (b) Gd-DTPA and (c) MION enhanced MRA maximum intensity projection images of the rat brain (scale bar = 5 mm).(d) Quantitative analysis of vascular imaging performance by calculating SNR values of inferior cerebral vein indicated by the red arrow in a before and after i.v.injection of P-UDIOC, Gd-DTPA, and MION (n = 3).Statistical analysis was performed using a Student's t-test, with *** indicating p < 0.001, ** indicating p < 0.01.(e) 3D volume image of the rat brain at 15 min postinjection of P-UDIOC, showing high spatial resolution of the vasculature (scale bar = 5 mm).The yellow arrows and labeled numbers in the image show the vascular details including 1, superior sagittal sinus; 2, transverse sinus; 3, anterior cerebral artery; 4, inferior cerebral vein; 5, basal vein; 6, middle meningeal artery.The inset image in (e) is an enlargement of the region indicated by the red dashed square (scale bar = 500 μm).Reproduced from [88].

Figure 3 .
Figure 3. Vascular imaging performance of P-UDIOC-enhanced UHF MRI.(a) P-UDIOC, (b) Gd-DTPA and (c) MION enhanced MRA maximum intensity projection images of the rat brain (scale bar = 5 mm).(d) Quantitative analysis of vascular imaging performance by calculating SNR values of inferior cerebral vein indicated by the red arrow in a before and after i.v.injection of P-UDIOC, Gd-DTPA, and MION (n = 3).Statistical analysis was performed using a Student's t-test, with *** indicating p < 0.001, ** indicating p < 0.01.(e) 3D volume image of the rat brain at 15 min post-injection of P-UDIOC, showing high spatial resolution of the vasculature (scale bar = 5 mm).The yellow arrows and labeled numbers in the image show the vascular details including 1, superior sagittal sinus; 2, transverse sinus; 3, anterior cerebral artery; 4, inferior cerebral vein; 5, basal vein; 6, middle meningeal artery.The inset image in (e) is an enlargement of the region indicated by the red dashed square (scale bar = 500 µm).Reproduced from [88].

Figure 4 .
Figure 4. Two spin phenomena between T2 and T1 contrast materials with different locations.(a) Left-the local magnetic field intensity of T1 contrast materials is reduced when located outside of the T2 contrast material.(b) Right-The local magnetic field strengths of T1 and T2 contrast materials are enhanced simultaneously when T1 contrast materials are located inside the T2 contrast materials.Reproduced from [103].

Figure 4 .
Figure 4. Two spin phenomena between T 2 and T 1 contrast materials with different locations.(a) Left-the local magnetic field intensity of T 1 contrast materials is reduced when located outside of the T 2 contrast material.(b) Right-The local magnetic field strengths of T 1 and T 2 contrast materials are enhanced simultaneously when T 1 contrast materials are located inside the T 2 contrast materials.Reproduced from [103].

Figure 5 .
Figure5.Schematic image of the core-shell-type dual-mode nanoparticle [MnFe2O4@SiO2@Gd2(CO3)2].The T1 contrast material is positioned on the shell to have direct contact with the water for high T1 contrast effects, and the superparamagnetic contrast material is located at the core, inducing a long-range magnetic field for the relaxation of water.Reproduced from[31].

Figure 5 .
Figure 5.Schematic image of the core-shell-type dual-mode nanoparticle [MnFe 2 O 4 @SiO 2 @Gd 2 (CO 3 ) 2 ].The T 1 contrast material is positioned on the shell to have direct contact with the water for high T 1 contrast effects, and the superparamagnetic contrast material is located at the core, inducing a long-range magnetic field for the relaxation of water.Reproduced from[31].

Figure 6 .
Figure 6.Schematic illustration for the synthesis of MnOx-SPION capped MSN and controlled drug release in response to the tumor microenvironment.In vivo T2 and T1 MRI images of a mouse upon injection of a MnOx-SPION@MSN solution through tail vein.Reproduced with permission from[111].Copyright from the American Chemical Society.

Figure 6 . 28 Figure 7 .
Figure 6.Schematic illustration for the synthesis of MnOx-SPION capped MSN and controlled drug release in response to the tumor microenvironment.In vivo T 2 and T 1 MRI images of a mouse upon injection of a MnOx-SPION@MSN solution through tail vein.Reproduced with permission from [111].Copyright from the American Chemical Society.Molecules 2024, 29, x FOR PEER REVIEW 15 of 28

Figure 8 .
Figure 8. Simultaneous T1 and T2 MRI imaging of liver and hepatic tumor (7 T), respectively.(a) T1w and (b) T2w in vivo MR images of BALB/c mice (top: coronal plane, bottom: transverse plane) before and after iv.injection of GdIO nanoparticles with a dose of 2.0 mg kg −1 .The regions of liver in the coronal planes were circled by dash lines.(c) T1w and (d) T2w in vivo MR images of nude mice orthotopically inoculated with HepG2 liver cancer cells (sagittal plane) before and after iv.injection of GdIO nanoparticles with a dose of 2.0 mg Fe kg −1 .Grey arrows: gallbladder, black arrows: liver, white dotted circles and white arrows: liver tumor.Reproduced from [103].

Figure 8 .
Figure 8. Simultaneous T 1 and T 2 MRI imaging of liver and hepatic tumor (7 T), respectively.(a) T 1w and (b) T 2w in vivo MR images of BALB/c mice (top: coronal plane, bottom: transverse plane) before and after iv.injection of GdIO nanoparticles with a dose of 2.0 mg kg −1 .The regions of liver in the coronal planes were circled by dash lines.(c) T 1w and (d) T 2w in vivo MR images of nude mice orthotopically inoculated with HepG2 liver cancer cells (sagittal plane) before and after iv.injection of GdIO nanoparticles with a dose of 2.0 mg Fe kg −1 .Grey arrows: gallbladder, black arrows: liver, white dotted circles and white arrows: liver tumor.Reproduced from [103].

Figure 9 .Figure 9 .
Figure 9. (a) Engineering the heterogeneous nanostructures for magnetic coupling of T1 and T2 contrast agents with (left) core/shell structures or (right) dumbbell structures.R0 is the radius of T2 contrast agent (iron oxide nanoparticles), R1 is the radius with the Gd shell, R2 is the radius of gold Figure 9. (a) Engineering the heterogeneous nanostructures for magnetic coupling of T 1 and T 2 contrast agents with (left) core/shell structures or (right) dumbbell structures.R 0 is the radius of T 2 contrast agent (iron oxide nanoparticles), R 1 is the radius with the Gd shell, R 2 is the radius of gold nanoparticles, and D represents the center-to-center distance between iron oxide and gold nanoparticles.(b) Design and characterization of dumbbell heterostructures for dual T 1 and T 2weighted MRI.Illustration of constructions of four different types of dumbbell-like or dumbbell heterostructures (Au_IONPs, HNTs, DB-HNTs, and extra-large DB-HNTs (XDB-HNTs)).Adapted from [127].

Fe 3 O 4 /
MnO hybrid nanocrystals were prepared based on seed-mediated growth of MnO on the surface of a Fe 3 O 4 , where the resulting structures depended on the size of the seed NP, producing core/shell spherical Fe 3 O 4 (5 nm)@MnO, dumbbell shaped Fe 3 O 4 (11 nm)/MnO, and flower-shaped Fe 3 O 4 (21 nm)/MnO hybrid NPs from 5 nm, 11 nm, and 21 nm Fe 3 O 4 NPs, respectively (Figure Molecules 2024, 29, x FOR PEER REVIEW 20 of 28 nanoparticles, and D represents the center-to-center distance between iron oxide and gold nanoparticles.(b) Design and characterization of dumbbell heterostructures for dual T1 and T2weighted MRI.Illustration of constructions of four different types of dumbbell-like or dumbbell heterostructures (Au_IONPs, HNTs, DB-HNTs, and extra-large DB-HNTs (XDB-HNTs)).Adapted from [127].

Figure 11 .
Figure 11.Serial T2-weighted MR images of the orthotopic xenograft model of a nude mouse liver for human HCC (arrows) before as well as 1 h, 6 h, 24 h and 72 h after the contrast injection of (a) Fe3O4/MnO dumbbell nanocrystals, (b) MnO nanocrystals, and (c) and Fe3O4 nanocrystals.Reproduced with permission from [109].Copyright from Elsevier.

Figure 11 .
Figure 11.Serial T 2 -weighted MR images of the orthotopic xenograft model of a nude mouse liver for human HCC (arrows) before as well as 1 h, 6 h, 24 h and 72 h after the contrast injection of (a) Fe 3 O 4 /MnO dumbbell nanocrystals, (b) MnO nanocrystals, and (c) and Fe 3 O 4 nanocrystals.Reproduced with permission from [109].Copyright from Elsevier.

Table 2 .
Main chemical and physical characteristics of paramagnetic ions used in MRI CAs.

Table 3 .
Summary of basic properties of the examples of DMCA NPs based on a typical T 1 contrast agent discussed in this section.Some of the systems in this Table have been used as theranostic agents.

Table 4 .
Summary of basic properties of the examples of NPs based on a typical T 2 ((U)SPION)) contrast agent discussed in this section.

Table 5 .
Summary of basic properties of the examples of NPs with a T 2 material inside a paramagnetic T 1 material.

Table 6 .
Summary of basic properties of NPs made of a paramagnetic T1 material inside a superparamagnetic T2 material or forming hybrid oligomers of different shapes (in bold).Their properties are compared with those of the NPs made of their components.