An update on clinical applications of magnetic nanoparticles for increasing the resolution of magnetic resonance imaging

Abstract

Today, technologies based on magnetic nanoparticles (MNPs) are regularly applied to biological systems with diagnostic or therapeutic aims. Nanoparticles made of the elements iron (Fe), gadolinium (Gd) or manganese (Mn) are generally used in many diagnostic applications performed under magnetic resonance imaging (MRI). Similar to molecular-based contrast agents, nanoparticles can be used to increase the resolution of imaging while offering well biocompatibility, poisonousness and biodistribution. Application of MNPs enhanced MRI sensitivity due to the accumulation of iron in the liver caused by discriminating action of the hepatobiliary system. The aim of this study is about the use, properties and advantages of MNPs in MRI.

Keywords:

• Diagnostic applications
• gadolinium
• magnetic nanoparticles
• magnetic resonance imaging

Introduction

Nanoparticles may provide advanced biomedical research tools based on polymeric or inorganic formulations. They can be used in many different medical and biological applications as in diagnostic test assays for early recognition of disease (Deerinck 2008, Gilmore and Spelke 2008, Hoffman-Amtenbrink et al. 2009, Son et al. 2007). The requirements for nanoparticles as biomedical agents span a broad range of narrative research strategies, materials and synthesis techniques. Diagnostic and therapeutic strategies need a different chemical or physical property of the particles included, which depend on the particular function played by the nanoparticles in that treatment (e.g. vector, porous receptacle, magnetic moment carrier). Magnetic nanoparticles (MNPs) are one category of these wide cancer-therapy designed nanoparticles (Goya et al. 2008, Oberdörster et al. 2005).

MNPs are spherical nanocrystals of 10–20 nm of size with a Fe³⁺ and Fe²⁺ core, sometimes enclosed by different molecules, such as dextran or polyethylene glycol (Lu et al. 2007, Pantic 2012, Sanvicens and Marco 2008). MNPs do not include any inert or non-degradable moieties, and it has been demonstrated that the iron oxide core of these agents is metabolized during two weeks, after which the free iron is used in the production of new red blood cells (Sosnovik 2007). The magnetic properties allow them to mark various biomolecules in new diagnostic techniques such as magnetic resonance imaging (MRI). In addition, they might be functionalized for dynamic targeting in vivo or for in vitro diagnostics (Lu et al. 2007, Pantic 2012, Sanvicens and Marco 2008).

The paradigmatic example is the MRI, a method which uses the magnetic moments of MNPs as a disorder of the proton resonance to achieve images (Goya et al. 2008). MRI is one of the high resolution (75–300 µs), most confident and multi-function imaging modalities of modern medicine. It also allows the exploring of cells, molecules and drug delivery carriers in the human body. The use of nanoparticles in modalities like MRI can significantly increase sensitivity, showing the potential for high-resolution molecular imaging. MRI has high spatial resolution (Ito and Cotsarelis 2005, Jun et al. 2007), which is compatible in nature, uses non-ionizing radiation and proposes multi-planar tomographic capabilities (Jun et al. 2007). Nanoparticles can be engineered to have magnetic characteristics that can be detected by MRI at low concentrations, and simultaneously include ligands which target particular molecules (Jun et al. 2007, Lodhia et al. 2012).

MRI scanners do not directly detect MNPs. They induce contrast improvement effects on the recreated MR images, either as a signal increase (positive contrast agents) or a signal decrease (negative contrast agents). MRI can supply a high spatial resolution as awaited, which is sometime of inadequate sensitivity by other hypointense areas. In attempting to find more complementary and efficient information about the physical structure and physiological function for disease diagnosis, novel nanoparticulate-based imaging platform is highly desired to combine the advantages of each imaging modality (Kolosnjaj-Tabi et al. 2014). In this review, we focus on the application of MNPs and their properties and advantages in MRI. We will also discuss the contrast agents (CAs) enhancement by using MNPs in this study.

MNPs in MRI

There are several imaging modalities clinically available such as MRI, computed tomography (CT), ultrasound (US), positron emission tomography (PET) or single photon emission computed tomography. Every technique has its own benefits and negatives but combining techniques can compensate for drawbacks in any specific technique (Table I) (Cheon and Lee 2008, Massoud and Gambhir 2003, Rudin and Weissleder 2003).

Table I. Comparison of several imaging modalities (Cheon and Lee 2008, Hachani et al. 2011, Massoud and Gambhir 2003).

	Source of imaging	Spatial resolution	Tissue penetrating depth	Sensitivity	Types of probe
MRI	Radiowave	25–100 mm	No limit	mM to µM (low)	Para-(Gd^3+) or superparamagnetic(Fe3O4) materials
PET	γ-ray	1–2 mm	No limit	pM (high)	Radionuclides (^18F, ^11C, ^13N, ^15O, ^124I, ^64 Cu)
CT	X-ray	50–200 mm	No limit	Not well characterized	High atomic number atoms (iodine, barium sulfate)

MRI is among today’s best studied imaging modalities for cellular tracking and labeling as it presents several advantages, for example, a great spatial (Kouhi et al. 2014, Smirnov 2008) and contemporary resolution, and the use of non-ionizing radiation. In medical performances, MRI detects the signal of the protons present in water, which will reply to a magnetic field in an environment-dependant method (i.e. different tissues) (Hachani et al. 2011).

The strength of the magnetic field in usual MRI scanners is in the range of 1–3 T, which is 20,000–60,000 times stronger than the Earth’s magnetic field. When a patient is introduced in the gantry, the hydrogen protons (¹H) align their spins along the direction of the magnetic field. The sum of each one of the magnetic moments of these spins represents the macroscopic magnetization vector (M) of the biological tissue (Lodhia et al. 2012).

Contemporary methods, containing US, CT and MRI, notice liver tumors only when they have grown to about 5 cm in diameter. By means of that time, the cancer is specifically aggressive, resisting chemotherapy and hard to remove surgically (Blasiak et al. 2013).

Lately, there have been many efforts to develop nano-sized difference agents for MRI and CT that are more active and practical than those present on today’s marketplace. Convinced scientists study superparamagnetic iron oxide particles (50–100 nm in size) as promising difference agents, because they have much greater magnetic susceptibility than traditional MR distinctions, for example, gadolinium (Gd) (Cuenca 2006, Pantic 2012). Contrarily to compute CT, a different functional imaging modality, it does not maintain by on ionizing radiation.

In 2009, it was anticipated that ultrasmall superparamagnetic elements of iron oxide (USPIO)-enhanced MRI in combination with diffusion-weighted magnetic character imaging could be a different, accurate and fast technique for detecting pelvic lymph node metastases even in normal-sized nodes of patients with bladder or prostate cancer (Abbasi et al. 2014, Pantic 2012).

Application of iron-based nanoparticles enhanced MRI sensitivity due to the accumulation of iron in the liver caused by discriminating action of the hepatobiliary system (Figure 1). This kind of contrast delivery does not apply, though, to most of the cancers consequently targeted, and active transfer is used (Blasiak et al. 2013).

Figure 1. An MRI (gradient echo, TR/TE = 100/4 ms, flip angle 30°, FOV = 5.5 × 2.5 cm, 256 × 256) of a mouse liver obtained at 9.4 T (A) before and (B) after injection of iron oxide (Nano-Ocean, Springdale, AR). The decrease of MR signal within the liver (yellow arrow) is visible (Blasiak et al. 2013).

To order a difference agent, the special of core and monolayer material is a critical step because this composition defines the primary physical and chemical properties and also reactivity, solubility and interfacial interactions. Most common core between the MRI CAs are paramagnetic lanthanide metals (Gd, manganese and dysprosium ion complexes) and superparamagnetic magnetite elements (iron oxides) (Yurt and Kazancı 2008). Iron oxide elements are widely studied in MRI applications as they change the relaxation times of tissues in which they are current and due to the low toxicity when associated to Gd chelates (Silva et al. 2012). From the opinion of view of MRI method, to increase MRI sympathy, two types of difference agents, providing positive or negative image difference, are used (Blasiak et al. 2013).

Contrast agents (CAs)

CAs are normally defined, created on their reduction properties, their magnetic properties and their biodistribution. When describing a CA based on reduction properties, the efficiency is described by the longitudinal and transverse relaxivity R₁ and R₂, respectively. The relativity reflects the alteration in the reduction rate as a function of the CA concentration. The relativities are affected by the size and the composition of these elements (Silva et al. 2012, Yurt and Kazancı 2008).

CAs have been progressive since the inception of MRI, to selectively alteration the longitudinal and transverse reduction times (T₁ and T₂) of ¹H in biological tissues. Such energy transmissions mostly happen though interactions taking place among the magnetic fundamentals and the spins of hydrogen protons in their vicinity (Caravan 2006, Caravan 1999). In parallel with the expansion of paramagnetic CAs, progress in the field of MRI CAs also occurred through the exceptional properties of iron oxide NPs. By opposition to paramagnetic molecules, which produce “positive” contrast in MRI, iron oxide NPs are well-known for their potential to rise the signal.

MRI positive CAs

CAs for longitudinal reduction of protons mostly comprise of paramagnets as these possess a great number of free electrons (lanthanides, for example) and shorten T₁, thus forming bright difference (Caravan 1999, Fallahzadeh et al. 2010). The shortening of T₁ depends solely on the direct conversation of energy among hydrogen protons in water molecules and the paramagnetic component’s outer shell electrons, hence an indirect correlation with the external of NPs (Hachani et al. 2011).

MRI negative CAs

Negative difference can be provided by superparamagnetic iron oxide NPs which condense the T₂ relaxation time and create a darker image. Two main factors perform a role in the contrast enhancement: the iron oxide core and its covering. The limitation of T₂ is determined by transverse reduction of hydrogen protons of water molecules surrounding the superparamagnetic iron oxide NPs (Gillis and Koenig 1987, Hachani et al. 2011).

MRI contrast enhancement

Both T₁ and $T_2^{*}$ can be condensed by the use of a magnetic CA. The most public CAs presently used are paramagnetic Gd ion complexes, though agents based on superparamagnetic nanoparticles are commercially presented, for example, “Feridex I. V.,” an iron oxide CA marketed by Advanced Magnetics Inc. (Rochester, IN) for the organ-specific targeting of liver lesions. The superparamagnetic elements used are magnetically saturated in the normal range of magnetic field strengths used in MRI scanners, thus founding a considerable locally perturbing dipolar field which leads, via Eq. (1(1) $\frac{1}{T_2^{*}}=\frac{1}{T_2}+\gamma \frac{\Delta B_0}{2}$(1) ), to a marked shortening of $T_2^{*}$ (Figure 2) along with a less marked reduction of T₁ (Elster 1994) (1) $\frac{1}{T_2^{*}}=\frac{1}{T_2}+\gamma \frac{\Delta B_0}{2}$(1) where ΔB₀ is the difference in the field brought about either through distortions in the homogeneity of the functional field itself, or by local differences in the magnetic susceptibility of the organization (Brown and Semelka 2015). $T_2^{*}$ is the smaller relaxation time and γ is the gyromagnetic ratio which γ = 2.67 × 10⁸rad s⁻¹ T⁻¹ for ¹H protons (Pankhurst et al. 2003).

Figure 2. Effect of magnetic particle internalization in cells on $T_2^{*}$ relaxation times: (A) the protons in cells tagged by magnetic particles have a shorter $T_2^{*}$ relaxation time than those in (B) untagged cells (Pankhurst et al. 2003, Yang et al. 2010).

Gd: Paramagnetic

Gadolinium-based difference agents are intravenous drugs used in diagnostic imaging actions to improve the quality of MRI or magnetic resonance angiography (MRA). Gd (III) containing MRI difference agents (often termed just “gado” or “gad”) are the most usually used for improvement of vessels in MRA or for brain tumor improvement associated with the degradation of the blood-brain barrier. For great vessels, for example, the aorta and its branches, the Gd (III) dose can be as low as 0.1 mmol per kg body mass. Higher attentions are frequently used for finer vasculature. Gd (III) chelates do not agreement the blood-brain barrier as they are hydrophilic. Therefore, these are convenient in enhancing lesions and tumors where the Gd (III) leaks out. In the rest of the body, the Gd (III) originally remains in the circulation but then dispenses into the interstitial space or is reduced by the kidneys (Hosseininasab et al. 2014). For instance, a free solubilized aqueous ion, Gd (III) is slightly toxic, but was usually viewed as safe when administered as a chelated complex. In animals at liberty Gd (III) ion displays a 100–200 mg/kg 50% lethal dose, but the LD50 is improved by a factor of 100 when Gd (III) is chelated, thus that its toxicity becomes comparable to iodinated X-ray difference complexes (Penfield and Reilly 2007). The chelating carrier molecule for Gd for MRI contrast use can be categorized by whether they are macrocyclic or have a right geometry and whether they are ionic or not. Cyclical ionic Gd (III) complexes are considered the least likely to release the Gd (III) ion, and, therefore, the safest.

Properties of MNPs for MRI

A very great fraction of molecular and cellular MRI submissions are based on the design and fabrication of MNPs consisting of: a) an inorganic nanocrystal core, for example, Fe₂O₃/Fe₃O₄, Gd₂O₃, MnO, NaGdF₄), b) a coating completed of small ligands or biocompatible biological molecules and c) an efficient anchor that is used to graft particular molecules, complementary imaging functionalities (e.g. radioactive atoms or fluorescent molecules), or medicinal complexes for drug delivery.

The MRI difference enhancement by MNPs is dependent on their structure, size, external things and degree of aggregation (Table II). It has been confirmed for instance that size is a critical parameter inducing the transverse relaxivity R₂ (Tromsdorf et al. 2007) although the surface contact influences the longitudinal relaxivity R₁ (Hachani et al. 2011, Tromsdorf et al. 2007). Great efforts have been devoted to construct multifunctional lanthanide-based nanoprobes for multi-modality imaging (Billotey et al. 2003, Corot 2006, Darrasse and Ginefri 2003, Karnoosh-Yamchi et al. 2014, Lévy et al. 1993, Pawelczyk et al. 2009, Smirnov 2008, Wilhelm et al. 2002). Some lanthanide ions with unpaired 4f electrons, like Gd³⁺ ions, are generally applied as CAs for MRI. Gd-, Yb- and Lu-doped nanostructures have been used as exceptional CT difference agents due to their stronger X-ray reduction and higher atomic number than those of usually used small iodinated molecules (Alizadeh et al. 2014, Chaudeurge et al. 2002, Fayol et al. 2002, Kolosnjaj-Tabi et al. 2014, Wilhelm et al. 2007).

Table II. Comparison of some nanoparticles with high transverse relaxivity for T₂ MRI (Benyettou et al. 2011, Hachani et al. 2011, Huang et al. 2011, Jung and Jacobs 1995, Ruiz et al. 2006).

Nanoparticle	Core size (nm)	Hydrodynamic diameter (nm)	R1 (mM^-1 s^-1)	R2 (mM^−1 s^−1)
Ferumoxide (feridex)  γFe2O3,Fe3O4–dextran	4.96	160	10.1	120
Fe3O4–PEG–(NH2)2	12	49	22 (1.5 T)	191 (1.5 T)
γFe2O3–alendronate	10	32.5	_	287 (4.7 T)
Fe3O4–casein	15	30	_	273 (3 T)

Iron oxide NPs are, in general, separated in two classes: small elements of iron oxide (SPIO) and ultrasmall components of iron oxide (USPIO). SPIOs are complete iron oxide cores of mean diameters in the range 3–20 nm; though the method agglomerates of hydrodynamic diameter usually superior to 50 nm. The theory of hydrodynamic diameter (Figure 3) mentions to the total actual diameter of an element suspended in a fluid and forming a colloid. The hydrodynamic diameter is mostly measured by means of laser analysis (dynamic light scattering), by evaluating the Brownian motion of the put off NPs. USPIOs instead, refer to the class of iron oxide NPs that are completed of iron oxide cores (3–20 nm diameter) bounded by a coating of molecules that preserve their individuality (mean diameter <50 nm).

Figure 3. Schematic representations of (A) the structure of functional magnetic nanoparticles (MNPs) for MRI applications and (B) a colloidal nanoparticle suspended in biological media.

Nanoparticle coating

NPs must form steady colloids in physiological media (blood, plasma, lymph and urine) for being considered in MRI applications. They must also show respectable biocompatibility and stretched vascular retention. The ligands and polymers, which are used as element coverings, must capably cover the particles and stimulate their individualization. NPs must be coated with a surface ligand or a polymer that either induces an electrostatic repulsion among the elements, or exert a strong steric attraction among them (Daraee et al. 2014). Amongst superparamagnetic difference agent, dextran-covered iron oxides are biocompatible and are excreted via the liver after the action. They are selectively taken up by the reticuloendothelial system, a system of cells lining blood vessels whose purpose is to remove foreign materials from the bloodstream; MRI contrast relies on the differential acceptance of different tissues (Lawaczeck et al. 1997, Pankhurst et al. 2003).

Conclusion

Superparamagnetic iron oxide NPs are proving to be a division of narrative probes practical for in vitro and in vivo cellular and molecular imaging. Superparamagnetic CAs have an improvement of producing an enhanced proton relaxation in MRI in comparison with paramagnetic ones. As a result, less amount of an SPIO mediator are required to dose the human body than a paramagnetic one. To relate the magnetic fluids to an MRI CA, an SPIO should be dispersed into a biocompatible and biodegradable carrier.

Acknowledgements

The authors thank the Department of Medical Nanotechnology, Faculty of Advanced Medical Science of Tabriz University for all the supports provided.

Declaration of interest

The authors declare that they have no competing interests. This work is funded by 2015 Drug Applied Research Center Tabriz University of Medical Sciences Grant.