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Nanotechnology Reviews

Editor-in-Chief: Hui, David

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Volume 3, Issue 3

Issues

Detection of magnetic nanomaterials in molecular imaging and diagnosis applications

Li Yao
  • Corresponding author
  • Beijing National Laboratory for Molecular Sciences, State Key Laboratory for Structural Chemistry of Unstable and Stable Species, Institute of Chemistry, Chinese Academy of Sciences, Beijing 100190, P.R. China
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/ Shoujun Xu
Published Online: 2014-03-22 | DOI: https://doi.org/10.1515/ntrev-2013-0044

Abstract

Advances in bionanotechnology promise to allow medical diagnosis and therapy through the channel of molecular imaging. Combining biological science and modern detection techniques, molecular imaging has the ability to penetrate biomedical processes at the molecular and cellular level. Magnetic nanoparticles (MNP), broadly defined as particles of tens of nm to approximately 2 μm in diameter in this review, are playing an increasingly important role in molecular imaging. They act as contrast agents to remarkably enhance the signal. The precise determination of the position and quantity of MNP is critical for these applications. This review describes the advances in the development of detection techniques for magnetic particles used in molecular imaging and diagnosis. The techniques are categorized as high magnetic field techniques and low magnetic field techniques. The high-field studies focus on magnetic resonance imaging (MRI). The ultra-low-field (ULF) studies include several of the most recent techniques: giant magnetoresistance sensors, superconducting quantum interference devices, atomic magnetometers, and magnetic particle imaging. The advantages and disadvantages of each method are discussed.

Keywords: atomic magnetometry; magnetic imaging; magnetic nanoparticles; molecular imaging

1 Introduction

Bionanotechnology is emerging as an important set of tools for scientists, engineers, and physicians to provide major advances in biological and medical sciences. The applications of nanomaterials in these fields are currently common, such as nanoparticles, quantum dots, nanorods, nanotubes, and other composites and nanostructured materials. Compared to the usual molecules, the nanomaterials are sufficiently large to have the advantage of realms of quantum behavior that are not otherwise accessible. Compared to the bulk materials, the nanomaterials show many unique characteristics, which are not presented in larger size, such as the exotic physical properties concerning the linear and nonlinear optical spectra [1], spin resonance spectra [2], temperature dependence of resistivity [3], and magnetic susceptibility measurements [4]. A number of remarkable phenomena have been observed that result in the in-depth knowledge of surface effect [5], nonlinear susceptibility enhancements [6], quantum tunneling [7], quantum phase transition [8], and quantum size-effect confinement [9].

Magnetic nanoparticles (MNP) exhibit excellent new characteristics such as superparamagnetism, high saturation field, high field irreversibility, extra anisotropy contributions. These excellent characteristics result from the finite size and surface effects that dominate the magnetic behavior of individual nanoparticles [10]. In 1930, Frenkel and Dorfman [11] predicted that a ferromagnetic particle with the size below 15 nm would produce a single magnetic domain. It means that this particle can hold a state of uniform magnetization at any field. When the measurement temperature is over a certain value, i.e., the blocking temperature, the magnetization behavior of these sufficiently small nanoparticles displays superparamagnetism. In an external magnetic field, they are magnetized similarly to a paramagnet, but possess extremely large susceptibilities. In the absence of external magnetic field, their magnetization appears to be on average zero.

A wide variety of applications have been demonstrated for MNP. For industrial applications, MNP are being used as catalysts or catalyst supports in chemistry [12], magnetic recording media in information technology [13], magnetic separation in waste water treatment [14], and magnetic seals in motors [15]. For biomedical applications, they are attractive for drug delivery systems, magnetic resonance contrast media, cancer therapeutic agents, magnetic immunoassay prober, and immune magnetic separation. Each potential application presents different properties of magnetic nanoparticles (MNP).

In biomedical applications, the particles are well dispersed usually in water as the so-called beads. Superparamagnetism provides much higher magnetization without the absence of the advantage of a stable colloidal suspension. Such particles commonly consist of the magnetically responsive component, such as iron, nickel, cobalt, and their chemical compounds. Superparamagnetic iron oxide (SPIO) particles such as maghemite (γ-Fe2O3) or magnetite (Fe3O4) are the most popular in biomedical application due to the good biocompatibility. Human tissues may contain iron or iron oxides in the form of hemosiderin, ferritin, and transferrin [16].

The human body uptake of particles depends on particle size including the iron core and the coating. Because the smallest capillaries in the body are 4 μm [17], larger particles will commonly deposit in the lungs [18]. The smaller particles are taken up through cells of the reticuloendothelial system. The venous sinuses of the spleen prefer to filter particles larger than 200 nm, while the small particles <100 nm will be phagocytosed through liver cells [19]. Under normal physiologic conditions, particles larger than 10 nm cannot penetrate the endothelium [20].

The surface modification with a biocompatible polymer coating is necessary to stabilize the nanoparticles in a biological suspension, and prevent the aggregation and immediate uptake by the reticulendothelial system. The polymer will also provide a functional group on the surface for a broad range of biomolecular binding. These MNP, formed through nanocrystalline synthesis, advanced polymer processing, or coating and functionalization strategies, have the potential to allow molecular targets to be imaged with different techniques, including many recent inventions and developments.

In this review, our aim is to give a comprehensive presentation of the recent state-of-the-art developments on the detection techniques for magnetic particles used in molecular and cellular imaging. The techniques are divided into two categories: high magnetic field, or simply high field (HF), techniques and ultra-low field (ULF) techniques. The HF techniques refer to techniques carried out in a magnetic field >1 T, which include magnetic resonance imaging (MRI) and its conjugation with other imaging modalities. The ULF techniques, which consist of the main part of this work, refer to techniques performed up to mT (10-3 T). We include the most recent developments for the following techniques: giant magnetoresistance (GMR) sensor, superconducting quantum interference devices (SQUID), atomic magnetometers (AM), and magnetic particle imaging (MPI).

2 High field studies

We will explain some fundamental principles of MRI first to highlight the differences between HF and ULF methods discussed in the next section. Readers can also get more complete information of conventional magnetic resonance imaging (MRI) in [21].

MRI can be thought of as occurring in three stages: polarization, encoding, and detection. First, a majority of nuclear spins of the object are polarized in an external magnetic field named Bp to point in one direction; second, a set of radiofrequency (RF) pulses are applied in a measuring magnetic field Bm to produce resonance absorption of the protons, during which gradient magnetic fields Gx,y,z are strategically applied to obtain spatial encoding; third, the nuclei return to equilibrium via relaxation and lose energy by emitting their own RF signal referred to as the free induction decay (FID) response signal. This measurement is reconstructed to obtain 3D MR images of the object. Here, we distinguish HF and ULF according to the intensity of measuring magnetic field Bm. In HF MRI, there is a single fixed field providing both Bp and Bm, typically by a large superconducting magnet of 1.5 T or higher to obtain high spatial resolution. In contrast, ULF techniques use different Bp and Bm, with the latter usually in the range of nT (10-9 T) to mT.

There are two principal relaxation processes termed T1 and T2. The longitudinal (or spin-lattice) relaxation time T1 and the transverse (or spin-spin) relaxation time T2 are the decay constants for the component of nuclear spin magnetization parallel and perpendicular to Bm, respectively. They are major intrinsic parameters that determine imaging contrast besides proton density. The T1 agents, commonly based on paramagnetic ions such as gadolinium (III) complexes, reduce the longitudinal relaxation time and increase the T1 signal intensity giving a positive contrast. There are four most important classical, clinically used, T1 agents: Gd-DOTA (France), Gd-DTPA (Germany), Gd-HPDO3A (Italy), and Gd-DTPA-BMA (USA) [22]. They are based on chelating agents with a branched or cyclic structure such as diethylenetriaminepentaacetic acid (DTPA) or 1,4,7,10-tetraazacyclododecane-N,N,N,N-tetraacetic acid (DOTA), respectively. SPIO nanoparticles are mostly used as contrast agents in MRI because of their significant negative enhancement effect on T2-weighted imaging pulse sequences. When placed in an external magnetic field, SPIO produce very high gradients that substantially disturb the local magnetic field to induce proton spin dephasing and consequently reduce the T2 relaxation time of the surrounding water. This predominant effect on the T2 relaxation time does not prevent the use of the properties of these agents on the T1 relaxation time when appropriate imaging sequences are chosen [23]. Because of the high anatomical spatial resolution of MRI, its potential for quantification, the low cytotoxicity of SPIO particles, and the promising recent progresses, magnetic particles now have a variety of applications in molecular and cellular imaging [24]. Molecular and cellular imaging gives the difference from traditional imaging for diagnostic application, in which probes known as biomarkers are used to help image particular targets or pathways. The purpose is to 1) better understand the fundamental molecular pathways inside organisms in a noninvasive manner, 2) diagnose the diseases such as cancer, and neurological and cardiovascular diseases at the molecular and cellular level, 3) improve the treatment of these disorders by optimizing the preclinical and clinical tests of new medication. It is difficult to strictly distinguish molecular imaging from cellular imaging. In this review, we divide the studies focused on the molecular targets into molecular imaging and the studies focused on the cellular targets into cellular imaging.

2.1 Molecular imaging

Molecular imaging usually can be considered as the noninvasive imaging of targeted biomolecules in biological processes of living organisms. For specific molecular imaging using MRI, many kinds of pharmacophores are conjugated to MNP, such as antibody/antigen, DNA/RNA, peptides, proteins, peptidomimetics, and small targeting ligands.

The earlier MRI studies showed that arabinogalactan-functionalized MNPs were able to identify the asialoglycoprotein receptor on hepatocytes [25, 26]. Asialofetuin-coated nanoparticles also presented the similar result of receptor-directed uptake [27]. Following the first observation of receptor-specific MRI, more works of molecular imaging are focused on the antibodies and peptides as specific pharmacophores for early diagnosis of diseases.

Immunoglobulins can be covalently linked to the nanoparticles. Most MNP surface is coated with dextran. The established method linking the amine groups of immunoglobulins to the alcohol groups of the dextran through a periodate-oxidation/borohydride-reduction reaction has been widely used [28, 29]. In 1991, human polyclonal immunoglobulin (Ig) G was attached to SPIO for receptor and antibody MRI [30]. When attached to antimyosin fragment antigen-binding (R11D10) and used for immune-specific MRI of cardiac infarcts, a marked decrease in the signal intensity of infracted myocardium was observed after intravenous administration in rats (100 μmol/kg) [31]. Intravenous injection of L6, a tumor-specific antibody conjugated to the nanoparticles showed a specific MRI pattern of enhancement of the tumors with the largest concentration of antibody in the area with the greatest density of tumor cells [32]. Similarly, immunospecific MRI using magnetite particles coated with monoclonal antibodies (mAbs) against epidermal growth factor receptors (EGFR) was useful in the diagnosis of squamous cell carcinoma of the esophagus [33]. Furthermore, some new versatile non-oxidative methods of attaching mAbs to magnetic particles were developed such as glutaraldehyde crosslinking [34], complexing through ultrasonication [35], biotin-streptavidin system [36, 37], and amine-sulfhydryl group linkage [38, 39]. Recently, an exchange encapsulation method was developed for MRI contrast enhancement and efficient cell targeting [40]. Through ligand exchange and subsequent encapsulation of MNP with casein, high-quality hydrophobic nanoparticles were transferred to the aqueous phase and exhibited substantially higher relaxivity and improved MRI contrast (Figure 1).

Casein-coated iron oxide nanoparticles (CNIO) for MRI contrast enhancement and efficient cell targeting. The transmission electron microscopy (TEM) images of (A) hydrophobic iron oxide (IO) nanoparticles before modification and (B) after being coated with casein (CNIOs). Inset of (A) is a photograph of hydrophobic IO nanoparticles dispersed in chloroform (left) and water-soluble CNIOs (right); inset of (B) is a negative-stained TEM image that shows a protein coating (bright layer surrounding the particles); (C) T2-weighted MR images of a mouse before and after injection of CNIOs at a dosage of 2.5 mg Fe/kg per mouse body weight. (Reproduced with permission from Ref. [40], Copyright ACS Publications, 2013.)
Figure 1

Casein-coated iron oxide nanoparticles (CNIO) for MRI contrast enhancement and efficient cell targeting. The transmission electron microscopy (TEM) images of (A) hydrophobic iron oxide (IO) nanoparticles before modification and (B) after being coated with casein (CNIOs). Inset of (A) is a photograph of hydrophobic IO nanoparticles dispersed in chloroform (left) and water-soluble CNIOs (right); inset of (B) is a negative-stained TEM image that shows a protein coating (bright layer surrounding the particles); (C) T2-weighted MR images of a mouse before and after injection of CNIOs at a dosage of 2.5 mg Fe/kg per mouse body weight. (Reproduced with permission from Ref. [40], Copyright ACS Publications, 2013.)

In the studies of peptides and proteins, transferrins are the important iron-binding blood plasma glycoproteins.

Human transferrin is encoded by the TF gene. When human holo-transferrin was covalently conjugated to the iron oxide nanoparticles (Tf-MION), transgene expression could be visualized directly in a live animal by MRI (Figure 2) [41]. The particles coated with short HIV-Tat peptides are able to efficiently internalize into hematopoietic and neural progenitor cells in quantities up to 10–30 pg of superparamagnetic iron per cell [42]. Surface modification of contrast agents with variable numbers of HIV-1 tat peptide demonstrates that higher numbers of tat peptide facilitate the cellular uptake of SPIO in a nonlinear fashion [43]. Conjugation of the C2 domain of synaptotagmin I to SPIO allowed noninvasive imaging of apoptotic cells both in vitro, with isolated apoptotic tumor cells, and in vivo, in a tumor treated with chemotherapeutic drugs [44]. The ultrasmall c(RGDyK) peptide-coated magnetic nanoparticels (<10 nm in hydrodynamic diameter) were prepared through 4-methylcatechol via the Mannich reaction. When administrated intravenously in mice, these particles accumulate preferentially in the integrin αvβ3-rich tumor area, which are readily tracked by MRI [45]. While the above studies provide examples of cases where molecular MR imaging using magnetic particles has proven to be successful, transposition to human applications is not yet available. The difficulty lies in the selection of the most effective pharmacophone for clinical imaging, which is limited by the low sensitivity of MRI and small biological targets.

Transgene expression in live animal by MR imaging. (A) Overexpression of engineered transferrin receptor (ETR) expression results in more cell uptake of the Tf-MION probe per hour than in control cells. The ETR cDNA sequence consists of the hTfR promoter, the coding sequence and the engineered 3 untranslated region (UTR) regulatory sequence. (B) In vivo MR imaging of a single mouse with stably expressing ETR (ETR+, left arrowheads) and nontransfected (ETR-, right arrowheads) flank tumors. T1-weighted coronal SE image (imaging time, 3.5 min; voxel resolution, 300×300×3000 μm). ETR- and ETR+ tumors have similar signal intensities. (C) T2-weighted gradient-echo image corresponding to the image in (A), showing substantial differences between ETR- and ETR+ tumors (imaging time, 8 min; voxel resolution, 300×300×3000 μm). As expected, ETR-mediated cellular accumulation of the superparamagnetic probe decreases signal intensity. These differences in MR signal intensity were most pronounced using T2- and T2*-weighted imaging pulse sequences, consistent with the increased transverse relaxation rate (R2) after cellular internalization. (D) Composite image of a T1-weighted spin-echo image obtained for anatomic detail with superimposed R2 changes after Tf-MION administration, as a color map. *Difference in R2 changes between the ETR+ and ETR- tumors. Scale bar (bottom left) represents 10 mm, n=1. (Reproduced with permission from Ref. [41], Copyright Nature Publishing Group, 2000.)
Figure 2

Transgene expression in live animal by MR imaging. (A) Overexpression of engineered transferrin receptor (ETR) expression results in more cell uptake of the Tf-MION probe per hour than in control cells. The ETR cDNA sequence consists of the hTfR promoter, the coding sequence and the engineered 3 untranslated region (UTR) regulatory sequence. (B) In vivo MR imaging of a single mouse with stably expressing ETR (ETR+, left arrowheads) and nontransfected (ETR-, right arrowheads) flank tumors. T1-weighted coronal SE image (imaging time, 3.5 min; voxel resolution, 300×300×3000 μm). ETR- and ETR+ tumors have similar signal intensities. (C) T2-weighted gradient-echo image corresponding to the image in (A), showing substantial differences between ETR- and ETR+ tumors (imaging time, 8 min; voxel resolution, 300×300×3000 μm). As expected, ETR-mediated cellular accumulation of the superparamagnetic probe decreases signal intensity. These differences in MR signal intensity were most pronounced using T2- and T2*-weighted imaging pulse sequences, consistent with the increased transverse relaxation rate (R2) after cellular internalization. (D) Composite image of a T1-weighted spin-echo image obtained for anatomic detail with superimposed R2 changes after Tf-MION administration, as a color map. *Difference in R2 changes between the ETR+ and ETR- tumors. Scale bar (bottom left) represents 10 mm, n=1. (Reproduced with permission from Ref. [41], Copyright Nature Publishing Group, 2000.)

2.2 Cellular imaging

Cellular MRI focuses on the visualization of the cell in its entirety instead of molecular structures expressed on the cell surface. One of the easiest and safest methods is spontaneous uptake of particles by phagocytic cells such as macrophages, microglia, and immature dendritic cells.

Cellular imaging with MNP was first used for hepatic imaging because removal of circulating particulates is accomplished in most species by macrophages resident in the liver and spleen. In the initial study of MNP evaluated as a contrast agent for MRI [46], doses ranging from 10 to 50 μmol/kg were administered intravenously to 15 patients. The clinical results appeared to confirm extensive preclinical data indicating that ferrite administered at a dose of 20 μmol/kg has the potential to significantly improve the performance of abdominal MR imaging. Based on the application of differentiation of normal and neoplastic liver tissues, the first SPIO product, ferumoxides (Feridex, Bayer HealthCare) was approved by the U.S. Food and Drug Administration (FDA) in 1996. Much smaller particles, monocrystalline iron oxide nanoparticles (MION) and ultrasmall superparamagnetic iron oxide (USPIO) were developed for MRI applications on lymph nodes (Figure 3) [47] and bone marrow [48], which have a longer blood half-life and are taken up by macrophages. In particular, the MRI with the macrophage-specific contrast agent USPIO could serve as a tool to predict disease severity. Using this approach, several disease models have been tested, including experimental autoimmune encephalomyelitis (EAE) [49–51], ischemia [52–54], allograft chronic rejection [55], lung allograft rejection [56], nephrotic syndrome [57], nephrotoxic nephritis [58], arthritis [59], and atherosclerotic lesions [60].

MRI and iron staining of an ipsilateral axillary lymph node 14 days after iron-labeled breast cancer cells were implanted into the thoracic mammary fat pad in CB17 SCID mice. (A) Cropped balanced steady-state free precession image of thorax region of mouse reveals a region of signal loss in the ipsilateral axillary node (arrow). (B) Section of lymph node stained with Perls’ Prussian Blue for iron detection at ×10 magnification shows iron-positive cells in the area of the node corresponding to the signal loss in the MR image (arrow). (C) Iron-positive area of the node at ×40 magnification, where strong staining can be seen (arrow). (Reproduced with permission from Ref. [47], Copyright Neoplasia Press, Inc, 2013.)
Figure 3

MRI and iron staining of an ipsilateral axillary lymph node 14 days after iron-labeled breast cancer cells were implanted into the thoracic mammary fat pad in CB17 SCID mice. (A) Cropped balanced steady-state free precession image of thorax region of mouse reveals a region of signal loss in the ipsilateral axillary node (arrow). (B) Section of lymph node stained with Perls’ Prussian Blue for iron detection at ×10 magnification shows iron-positive cells in the area of the node corresponding to the signal loss in the MR image (arrow). (C) Iron-positive area of the node at ×40 magnification, where strong staining can be seen (arrow). (Reproduced with permission from Ref. [47], Copyright Neoplasia Press, Inc, 2013.)

For nonphagocytic cells that do not spontaneously imbibe nanoparticles, surface modification of MNP have to adapt for intracellular uptake and magnetic labeling. One method for intracellular particles is to use the internalizing monoclonal antibodies (mAbs) [61–64]. It is so species-specific that each new antibody needs to be synthesized for each different animal research. This method is thus not highly compatible with the utilization of a xenogeneic protein in the clinical use. A widely used strategy is anionic magnetic nanoparticles. They are able to efficiently internalize into macrophages where they concentrate within micrometric endosomes, conferring on them a high magnetic susceptibility [65]. The analysis of particle uptake kinetics indicates that the nonspecific interaction of negatively charged SPIO with macrophages was a two-step process: binding onto the cell surface (a Langmuir adsorption) and cell internalization (a saturable mechanism) [66].

Owing to excellent spatial resolution and whole-body soft-tissue contrast, MRI is suitable for monitoring the migration and tracking of magnetic-labeled stem cell. One main use of stem cells in medicine is as a source of donor cells to be used as therapy to replace damaged or missing cells and organs. Stem cells are also useful for creating models of human disease and for drug discovery. In some cases, in vitro magnetic labeling is not sufficient. The cationic transfection agents such as polylysine, protamine sulfate, lipofectamin, and dendrimer are helpful for the internalization of magnetic particles [67–72]. They are conjugated to the anionic ferumoxide magnetic particles through electrostatic interactions. The requirement of transfection agent to improve the internalization rate depends on the cell type and particle size. SPIO particles with a diameter of about 100–150 nm have been shown to be more efficient for cell targeting than USPIO nanoparticles with diameters of 20–40 nm [73]. It may be due to the additional phagocytic and endocytotic uptake of nanoparticles with bigger size. Rabbit tendon stem cells (TSCs) labeled by incubation with 50 μg/ml SPIO could be detected by MRI both in vitro and in vivo. The labeling efficiency of TSCs reached as high as 98% [74]. The first clinical MRI cell tracking was performed in the Netherlands to examine the use of ferumoxide-labeled dendritic cells [75]. The second clinical MRI study was preformed in China. The autologous neural stem cells were labeled with ferumoxides and stereotactically injected near the area of patient’s brain injury. MR images were obtained at 3 T. The signal intensity had disappeared completely 7 weeks after injection [76]. Because of the emergence of stem cell therapy and the need for high-resolution noninvasive tracking methods of clinical translation, MRI cell tracking has become a robust method.

Other methods of magnetic labeling have also been described but are not as widely used. One approach is electroporation. It uses an electrical pulse to induce a change in the electrochemical permeability of the cell membrane and enables efficient intracytoplasmic labeling of cells [77]. This approach is commonly used for transferring DNA and chemotherapeutic drugs into cells and was recently applied to stem cell labeling with MNP [78]. MRI tracking of adipose-derived mesenchymal stem cells (MSC) was still visible 14 days after injection (Figure 4).

Magnetic nanoparticles as positive T1 contrast agents for labeling and MRI tracking of adipose-derived mesenchymal stem cells (MSC). (A) Labeling of MSC. (B) In vivo MRI of transplanted with unlabeled MSCs. (C) In vivo MRI of transplanted with HMnO@mSiO2-labeled MSCs and were still visible 14 days after injection. (Reproduced with permission from Ref. [78], Copyright ACS Publications, 2011.)
Figure 4

Magnetic nanoparticles as positive T1 contrast agents for labeling and MRI tracking of adipose-derived mesenchymal stem cells (MSC). (A) Labeling of MSC. (B) In vivo MRI of transplanted with unlabeled MSCs. (C) In vivo MRI of transplanted with HMnO@mSiO2-labeled MSCs and were still visible 14 days after injection. (Reproduced with permission from Ref. [78], Copyright ACS Publications, 2011.)

2.3 Multimodal approach

Nanoparticles compose an important molecular and cellular imaging agent class for multimodality preclinical imaging in addition to MRI, such as positron emission tomography (PET), computed tomography (CT), single photo emission CT (SPECT), ultrasound, optical imaging, due to their properties in fluorescence, bioluminescence, and optical coherence. MRI offers the advantage of high spatial resolution, but suffers from poor sensitivity and high cost. PET is supersensitive with a relatively low spatial resolution. Fluorescence imaging is extremely sensitive, but lacks the ability to provide anatomical resolution. Therefore, multifunctional magnetic nanoparticles with the advantages of two or more imaging modalities are becoming the new attractive forms in molecular and cellular imaging. The imaging modalities with high sensitivity are commonly combined with the others, which have high spatial resolution. With the development of the adapted techniques combing two or more types of analysis in a single machine, multimodal magnetic nanoparticles are going on growing for the applications of early stage cancer diagnosis, guided stem cell therapies, drug delivery, pathogen detection, and gene therapy.

Many polymeric coating materials have been used for multimodal imaging, such as dextran, carboxymethylated dextran, carboxydextran, starch, polyethylene glycol (PEG), arabinogalactan, glycosaminoglycan, and organic siloxane [79]. They provide a large quantity of carboxylic or amine groups to label nanoparticles with affinity ligands, radioactive isotopes, or fluorochromes for multimodal detection (Figure 5) [80].

Concept of a multimodal iron oxide nanoparticle. (A) Iron oxide core is engulfed by a polymer shell (e.g., dextran). Amine groups bind affinity ligands and beacons for various imaging modalities (B–D). (B) Surface derivatization with radionuclides promotes nuclear imaging, e.g., positron emission tomography (PET, nanoparticles targeting tumor-associated macrophages) or single-photon emission computed tomography (SPECT, targeting activated platelets). (C) Superparamagnetic iron oxide core causes signal decrease in T2*-weighted MRI [arrows show atherosclerotic plaque in apolipoprotein E-deficient (apoE-/-) mouse aorta]. (D) Fluorochromes attached to the nanoparticle surface enable optical imaging, such as fluorescence molecular tomography (FMT), fluorescence reflectance imaging (FRI) or florescence microscopy (FM). (Reproduced with permission from Ref. [80], Copyright John Wiley and Sons, 2013.)
Figure 5

Concept of a multimodal iron oxide nanoparticle. (A) Iron oxide core is engulfed by a polymer shell (e.g., dextran). Amine groups bind affinity ligands and beacons for various imaging modalities (B–D). (B) Surface derivatization with radionuclides promotes nuclear imaging, e.g., positron emission tomography (PET, nanoparticles targeting tumor-associated macrophages) or single-photon emission computed tomography (SPECT, targeting activated platelets). (C) Superparamagnetic iron oxide core causes signal decrease in T2*-weighted MRI [arrows show atherosclerotic plaque in apolipoprotein E-deficient (apoE-/-) mouse aorta]. (D) Fluorochromes attached to the nanoparticle surface enable optical imaging, such as fluorescence molecular tomography (FMT), fluorescence reflectance imaging (FRI) or florescence microscopy (FM). (Reproduced with permission from Ref. [80], Copyright John Wiley and Sons, 2013.)

The efficient delivery of nanomaterials to specific targets for in vivo biomedical imaging is hindered by rapid sequestration by the reticuloendothelial system (RES) and consequent short circulation times. To overcome these problems, a new stealth PEG polymer containing terminal radio-labeled bisphosphonate (BP) group was conjugated to the surface of USPIO for visualization with MRI and SPECT. It affords a dual-modality contrast with low RES uptake and long blood circulation times [81]. The unique multimodal nanoparticles with high density and lipoprotein-mimicking properties were developed by incorporation of gold, iron oxide, or quantum dot nanocrystals for computed tomography, magnetic resonance, and fluorescence imaging, respectively [82]. When internalized into phagocytes, they can be used to efficiently detect macrophage presence in atherosclerotic plaques. The ability to image cardiomyocyte (CM) apoptosis in heart failure could facilitate more accurate diagnostics and optimize targeted therapeutics. The uptake of annexin-labeled magnetic nanoparticles (AnxCLIO-Cy5.5) occurred early by CM apoptosis after myocardial infarction. The specificity of imaging signal was detected with high sensitive T2*-weighted MRI and colocalization ex vivo fluorescence imaging [83]. In tumor imaging, the triple functional probe was developed for MRI/PET/near-infrared fluorescence (NIRF). The SPIO coated with human serum albumin (HSA) were dually labeled with 64Cu-DOTA and Cy5.5 and tested in a subcutaneous U87MG xenograft mouse model. With the compact HSA coating, the HSA-IONPs showed massive accumulation in lesions, high extravasation rate, and low uptake of the particles by macrophages at the tumor area [84]. The combination of different imaging modalities connects noninvasively with potentially more specific invasively gained data and thus facilitates insight into complex basic biology.

2.4 Limitations of HF MRI

Usually for MRI, the higher the magnetic field, the better signal-to-noise ratio, higher contrast-to-noise ratios, and higher spatial resolution can be achieved. However, there are technological, physical, and safety limitations that restrict the full realization of the high-field benefits. For example, there is a limit of the magnetic field in the range of 10–12 T because of the limits on the superconducting materials. Magnetic field inhomogeneity is another major drawback in high-field MRI [85]. The non-uniformity causes regions with increased and decreased signal intensities in the images. Moreover, the increase in field strength causes the increase in equipment weight and energy consumption, which make MRI bulky and expensive.

There are also potentially safety issues involving the possibility of ferromagnetic projectiles. Patients with metal implants should not be able to have an MRI scan. Besides, the FDA specified in 2003 that neonates younger than 1 month have not been ruled out above 4 T, nor ruled out above 8 T for older children and adults. Although there have been no health hazards reported with physiological exposure to higher fields, it is known that the incidence of the magnetohydrodynamic effects and human response increases with the magnetic field strength. Nausea, vertigo, headache, tingling, numbness, visual disturbances (phosphenes), and discomfort associated with tooth fillings have been reported in patients moving in high fields [86–88].

3 Ultra-low field studies

HF MRI detects the presence of MNP, indirectly by detecting the water surrounding the magnetic particles. However, ULF techniques are able to directly detect the magnetization from magnetic particles because of the low magnetic background from biological samples, which enables the high sensitive identification. Here, we cover four types of ULF techniques for molecular magnetic imaging and analysis employing magnetic nanoparticles: giant magnetoresistance (GMR) sensor, superconducting quantum interference devices (SQUID), atomic magnetometers (AM), magnetic particle imaging (MPI). Each type of ULF technique has its own advantages and limitations, which include resolution, sensitivity, cost, and speed (Table 1).

Table 1

List of ultra-low field techniques for molecular and cellular magnetic imaging and analysis.

3.1 Giant magnetoresistance (GMR) sensor

GMR sensors consist of thin-film structures made from alternating magnetic and nonmagnetic layers. In 1988, the giant magnetoresistance effect was discovered [89]. The electrical resistance of this film undergoes a significant change as a function of the applied external magnetic field strength. The practical significance of this experimental discovery was recognized by the Nobel Prize in Physics awarded to Fert and Grünberg in 2007. This development has an immeasurable impact on magnetic data storage technology [90]. The successes of GMR spin valve (SV) and magnetic tunnel junction (MTJ) sensors in hard disk drives and magnetic memories offer an inspiration for their use in magnetic biosensors, a growing field with great promise and worldwide participation. GMR biosensors operate in a similar fashion to the enzyme-linked immunosorbent assay (ELISA) where a capture antibody is immobilized on the surface of the sensor, and the target of interest selectively binds to the antibody. The assay is completed by introducing a detection antibody labeled with an externally observable tag. Unlike ELISA where the tag is typically a fluorescent molecule, GMR biosensors rely on magnetic particles. When a magnetic particle is in the vicinity of the sensor, the magnetic field on the sensor will be slightly altered, and this can be detected. Usually, a vertical magnetic field (<100 mT) generated by a Helmholtz coil was applied as induced field in the detection of magnetic particles. The sensors are coated with an ultrathin but corrosion-resistive passivation layer, allowing the bimolecular targets and tags to be very close to the magnetic sensing layer. A group at the Naval Research laboratory and NVE Corporation in the USA first demonstrated a magnetic biosensor system, which they called BARC (Bead Array Counter) [91]. Several groups have continued the research and development of GMR biosensing technology [92–96]. The commercially available magnetic particles used by most of these groups have a size in the range from 0.1 μm to 3 μm such as Dynal M280 (2.8 μm), Bangs CM01N (350 nm), Ademtech (300 nm), Nanomag-D (250 nm), Miltenyi MACS (40 nm).

Based on the BARC technique, the threshold for detection of M280 is approximately 10 beads per 200-μm diameter sensor [97]. Streptavidin-coated magnetic beads were applied to spin valve sensors for the sensitive and specific detection of the marker protein S100ββ at 27 pg/ml level of sensitivity [98]. By using 50 nm MACS magnetic nanoparticles, cancer marker was detected by SV sensors in 50% serum at sub picomolar concentrations with a dynamic range of more than four decades. With the addition of nanotag amplification, the sensitivity extended into the low fM (10-15 m) concentration range [99]. A highly sensitive magnetoresistive biochip was successfully applied to the detection of oligonucletide hybridization events. The biological detection limit of the device was significantly improved by three orders of magnitude, from 1 pm to 1 fm, by the use of an efficient magnetic focusing system associated to prehybridization-labeled samples [100]. Most GMR biosensing platforms employ a sandwich approach. The three-layer competition-based assay could offer ultrasensitive detection and quantification (Figure 6). This method was used to detect as few as 1000 copies of endoglin at concentrations as low as 83 fm with high detection specificity from unprocessed human urine samples [101].

(A) Two-layer competition detection scheme. The GMR sensors were first functionalized with capture antibodies (a1). Nanoparticle-labeled endoglin and unlabeled endoglin were then mixed and applied on the sensor surface to compete for capture antibodies (a2). (B) Three-layer competition detection scheme. The GMR sensors were first functionalized with capture antibodies. Biotin-labeled endoglin and unlabeled endoglin were then mixed and applied on the sensor surface to compete for capture antibodies (b1). Subsequently, streptavidin-FeCo nanoparticle conjugates were applied for binding (b2 and b3). (Reproduced with permission from Ref. [101], Copyright ACS Publications, 2011.)
Figure 6

(A) Two-layer competition detection scheme. The GMR sensors were first functionalized with capture antibodies (a1). Nanoparticle-labeled endoglin and unlabeled endoglin were then mixed and applied on the sensor surface to compete for capture antibodies (a2). (B) Three-layer competition detection scheme. The GMR sensors were first functionalized with capture antibodies. Biotin-labeled endoglin and unlabeled endoglin were then mixed and applied on the sensor surface to compete for capture antibodies (b1). Subsequently, streptavidin-FeCo nanoparticle conjugates were applied for binding (b2 and b3). (Reproduced with permission from Ref. [101], Copyright ACS Publications, 2011.)

Recently, novel high magnetic moment nanoparticles have been prepared to replace the larger particles for GMR detection. The newly developed synthetic anti-ferromagnetic-layered nanoparticles allowed DNA detection with high sensitivity (10 pm) and tunable responses to a small external magnetic field gradient [102].

A detecting system based on a GMR sensor and 12.8-nm high-moment cubic FeCo nanoparticles in a 1:1 ratio was used to linearly detect 600–4500 copies of streptavidin. This sensitivity is expected to suffice for detection of all known potential biomarkers from body fluid samples of 10 nl or less [103].

GMR sensors can be arrayed and multiplexed to perform analysis on a panel of proteins or nucleic acids in a single complementary metal oxide semiconductor (CMOS) chip, yielding a low-cost system for future biological diagnostics. A circuit architecture was built scalable for larger sensor arrays of 64 individually addressable sensors while maintaining a high readout rate scanning the entire array in <4 s. The results showed the ability to distinguish a multiplicity of proteins present at a wide variety of concentrations in a single reaction well [104]. By using magnetic particles as labels to detect DNA hybridization, a prototype device was illustrated, which is a tabletop unit containing electronics, a chip carrier with a microfluidic flow cell, and a compact electromagnet. It is capable of simultaneous detection of eight different analytes in a microfabricated chip [105].

Recently, the GMR platform has been scaled to over 100,000 sensors per cm2, which can be used to measure the binding kinetics of antibody-antigen binding as low as 20 zeptomoles (10-21) of solute (Figure 7). The real-time binding assay and kinetic model has been extended to visualize protein-binding events in both space (due to the high density of the array architecture) and time (due to the rapid and real-time readout) [106]. The potential clinical applications of this method are vast ranging from investigating drug on-target and off-target cross-reaction binding kinetics, to in vitro clinical assay development, to targeted molecular imaging for early cancer diagnostics.

GMR nanosensor system for kinetic analysis. (A) optical micrograph showing the GMR sensor architecture comprising 72 stripes connected in parallel and in series. Inset: SEM image of one stripe of the GMR sensor with several bound magnetic nanoparticle tags. (C) Schematic representation of a magnetically labeled antibody, drawn to scale. The magnetic tag comprises a dozen iron-oxide cores embedded in a dextran polymer and then functionalized with antibody or receptor. (B) Schematic depicting the GMR sensor array functionalized with monoclonal anticarcinoembryonic antigen (CEA) capture antibody (not to scale). The solution above the sensor array is composed of magnetically labeled anti-CEA detection antibodies. The schematic includes a pipette tip containing a solution of CEA protein before injection. (C) Once the CEA antigen is introduced into the solution above the sensor array, radial transport of CEA antigen from near the centre of the array is monitored in real time. Magnetically labeled detection antibodies capture CEA protein and bind to anti-CEA antibodies on the GMR sensor surface to form detectable sandwich structures. (D) Visualization of CEA protein surface concentration at different times using a high-density GMR sensor array. The units of the y-axis are presented in changes in MR normalized to the initial MR in ppm. (Reproduced with permission from Ref. [106], Copyright Nature Publishing Group, 2011.)
Figure 7

GMR nanosensor system for kinetic analysis. (A) optical micrograph showing the GMR sensor architecture comprising 72 stripes connected in parallel and in series. Inset: SEM image of one stripe of the GMR sensor with several bound magnetic nanoparticle tags. (C) Schematic representation of a magnetically labeled antibody, drawn to scale. The magnetic tag comprises a dozen iron-oxide cores embedded in a dextran polymer and then functionalized with antibody or receptor. (B) Schematic depicting the GMR sensor array functionalized with monoclonal anticarcinoembryonic antigen (CEA) capture antibody (not to scale). The solution above the sensor array is composed of magnetically labeled anti-CEA detection antibodies. The schematic includes a pipette tip containing a solution of CEA protein before injection. (C) Once the CEA antigen is introduced into the solution above the sensor array, radial transport of CEA antigen from near the centre of the array is monitored in real time. Magnetically labeled detection antibodies capture CEA protein and bind to anti-CEA antibodies on the GMR sensor surface to form detectable sandwich structures. (D) Visualization of CEA protein surface concentration at different times using a high-density GMR sensor array. The units of the y-axis are presented in changes in MR normalized to the initial MR in ppm. (Reproduced with permission from Ref. [106], Copyright Nature Publishing Group, 2011.)

GMR sensors have the advantage of small size, light weight, low cost and power consumption, which make it useful as biochip for molecular analysis. But its sensitivity is relatively poor compared to other magnetometers such as SQUID and AM. Because the magnetic signal of the MNP decreases rapidly when the distance between the particles and the sensor increases, the magnetic particles to be detected are limited to stay on the sensor surface, and GMR sensors cannot be applied for in vivo imaging.

3.2 Superconducting quantum interference devices (SQUID)

SQUID is recognized as the most sensitive magnetometer, achieving a noise level of the order of several fT/Hz1/2 [107]. It is so sensitive that a field change of 5 aT (5×10-18 T) is detectable within a few days of averaged measurements [108]. The range of application of SQUID is very wide, such as routine measurement of condensed-matter magnetic susceptibility, gravitational wave detection, nuclear magnetic resonance (NMR) and MRI detection, studies of palaeomagnetism, nondestructive testing, and underwater ordinance detection [109–115]. However, the most important application has been in the area of biomagnetism, the detection of the weak magnetic fields produced by the human brain, heart, and other organs [116, 117].

The operation principles of SQUID are based on two effects observable only in the presence of superconductivity: the magnetic flux quantization and the Josephson effect [118–120]. There are two main types of SQUID. One is direct current (DC) SQUID consisting of a superconducting ring interrupted by two Josephson junctions. The other is radio frequency (RF) SQUID working with only one Josephson junction. The current induced in the SQUID ring by an external magnetic flux results in a change of voltage across the junction. This amplified voltage forms the SQUID output signal.

SQUID are used in ULF MRI with microtesla-range magnetic fields for the image that has been pioneered by John Clarke, Alexander Pines, and coworkers [121, 122]. The low signal-to-noise ratio is partially compensated for by pre-polarizing the sample in a field up to 200 mT and using SQUID for signal detection. The advantages of ULF MRI lie in unique low-field contrast mechanisms, flexibility in the sequence design, and the possibility to construct a silent scanner with an open geometry. It is also compatible with magnetoencephalography (MEG), which uses SQUIDs to record the magnetic field produced by neuronal activity. With a hybrid scanner combining MEG and MRI, both the structure and function of the human brain can be studied with a single device [123–126]. A new type of MNP-based imaging method was evaluated through standard clinical MEG systems equipped with SQUID. The results indicate that the magnetic signals of nanoparticles can be detected without magnetization using standard clinical MEG, and only bound particles produce detectable signals [127]. SQUID have also been used to track the naturally occurring magnetic particles in human organs, for example, examination of human liver iron stores [128] and iron accumulation in lungs [129].

We are mainly concerned on synthetic MNPs that are used as labels for molecular and cellular imaging and detected by SQUID. As having long been the most sensitive magnetometer until recent developments in AM, SQUID stands out as a promising technique for direct mapping of MNP. Pseudo lymph nodes containing USPIO were made and tested for breast cancer detection. At the distance of 20 mm, the detectable minimum weight of the iron was 40 mg [130]. Magnetic signal from a 1-cm3 tumor loaded with MNP (LCDIO) was successfully measured by the SQUID scanner at a depth of 7 cm into the body [131]. Figure 8 presents the co-registration of SQUID images and standard 4.7 T MRI of mice for breast cancer detection, resulting in good agreement. Single-core iron oxide nanoparticles were conjugated to a Her2 monoclonal antibody and targeted to Her2-expressing MCF7/Her2-18 (breast cancer cells) in nude mouse. The SQUID detection limit of 125,000 nanoparticle-labeled cells was obtained at 3 cm from the sensors [132]. In addition, field-dependent T1 relaxation and imaging of ferrofluids were reported. It combines magnetic fluids consisted of dextran-coated MNP and SQUID-detected NMR and imaging, which shows the potential to image tumors [133]. Based on magnetic relaxation of MNPs, magnetic imaging of the tracers was obtained directly by mapping the magnetization decays. The method relies on the principle that Néel relaxation of the magnetic nanoparticles is faster in a finite magnetic field than in the absence of the field [134].

Imaging of a nude mouse with two xenograft MCF7/Her2–18 tumors injected with Her2-conjugated magnetic nanoparticles. (A) Photograph of the mouse on the relaxometry imaging stage. (B) T2-weighted MRI after intratumoral injection of nanoparticles  (normal gray scale). (C) SQUID-detected magnetic relaxometry confidence intervals are centered at the positions of the detected dipole sources. The size of the confidence intervals indicates the uncertainty in the determination of the source position in the X and Y directions. (D) Co-registry of the relaxometry confidence intervals and the MRI (reverse gray scale). The MRI (4 cm FOV) was scaled to the correct size relative to the relaxometry coordinate grid. The MRI was then translated to the same origin used for the relaxometry measurement (X=0 line coincides with the spine, Y=0 line bisects tumors). (Reproduced with permission from Ref. [132], Copyright John Wiley and Sons, 2012.)
Figure 8

Imaging of a nude mouse with two xenograft MCF7/Her2–18 tumors injected with Her2-conjugated magnetic nanoparticles. (A) Photograph of the mouse on the relaxometry imaging stage. (B) T2-weighted MRI after intratumoral injection of nanoparticles (normal gray scale). (C) SQUID-detected magnetic relaxometry confidence intervals are centered at the positions of the detected dipole sources. The size of the confidence intervals indicates the uncertainty in the determination of the source position in the X and Y directions. (D) Co-registry of the relaxometry confidence intervals and the MRI (reverse gray scale). The MRI (4 cm FOV) was scaled to the correct size relative to the relaxometry coordinate grid. The MRI was then translated to the same origin used for the relaxometry measurement (X=0 line coincides with the spine, Y=0 line bisects tumors). (Reproduced with permission from Ref. [132], Copyright John Wiley and Sons, 2012.)

The application of SQUID for detecting targeted MNP was also attempted on molecular and cellular immunoassays. Based on the magnetization behavior and mechanism of MNP, there are three basic methods for immunoassays: relaxation, remanence, and susceptibility. Magnetic particles undergo two relaxation mechanisms after the external magnetic field is removed, Brownian relaxation and Néel relaxation [135]. The effective relaxation rate is determined by the sum of the Brownian relaxation rate and the Néel relaxation rate. The relaxation method is based on the measurement of relaxation rate. When adding 50 nm of MNP coated with antibody to an aqueous sample containing bacteria (Listeria monocytogenes), the unbound particles randomized by Brownian relaxation are too quick to be detected. In contrast, the particles bound to bacteria relax in about 1 s. This Néel relaxation process detected by SQUID, exhibited a limit of detection of 1.1×105 bacteria in a 20-μl sample volume [136]. This technique also has been demonstrated with a model system of liposomes carrying the flag epitope, which requires no more than 5×104 MNP to register a reproducible signal [137]. In the remanence method, Brownian relaxation progresses in micro seconds. But Néel relaxation can take a few hours, which is sufficiently long for SQUID to measure the magnetization of magnetic particles bound to the target. The typical difference of the remanence method is the requirement of sample scanning to produce a magnetic signal change for SQUID detection. This method was able to detect 25 nm of Fe2O3 nanoparticles with the sensitivity of 10 ng at a distance of 1.7 cm [138]. A DNA hybridization assay was investigated on sample coverslips. After washing off unbound magnetic particles, the surface was scanned by SQUID to detect hybridization [139]. In the immunoassay of antibody-antigen, the remanence measurement was specific only for bound components, eliminating the washing step [140]. The detection limit of IgE could reach 15 attomol [141]. For detecting multiple samples, a prototype of the magnetic immunoassay system was built with a 12-cell disk, which could detect 30 pg of 25-nm Fe3O4 nanoparticles [142].

In recent susceptibility studies, a new method referred to as SQUID-based immunomagnetic reduction (IMR) was developed for immunoassays [143]. IMR detect the mixed-frequency AC magnetic susceptibility of magnetic particles. When a portion of the particles conjugate with the target molecules, the average size of particles increases, which results in lower AC magnetic susceptibility of particles. The detection throughput increased using the multi-channel IMR [144, 145]. It was demonstrated that the detection limit is approximately 10 pg/ml for the malignancy biomarker, vascular endothelial growth factor (VEGF), in a serum sample (Figure 9) [146].

In vitro assay of vascular endothelial growth factor (VEGF) using immunomagnetic reduction (IMR). (A) Under external multiple ac magnetic fields, magnetic nanoparticles oscillate with the multiple ac magnetic fields via magnetic interaction. Thus, the reagent under external multiple ac magnetic fields shows a magnetic property, called mixed-frequency ac magnetic susceptibility. (B) Via the antibodies on the outmost shell, magnetic nanoparticles associate with magnetically label targets. With the association, magnetic nanoparticles become larger or clustered. The response of these larger magnetic nanoparticles to external multiple ac magnetic fields becomes much less than that of originally individual magnetic nanoparticles. (C) Detected VEGF concentrations in serum using SQUID-based IMR for normal people and patients with hepatocellular carcinoma (HCH) or colorectal cancer (CRC). (Reproduced with permission from Ref. [146], Copyright Hindawi Publishing Corporation, 2013.)
Figure 9

In vitro assay of vascular endothelial growth factor (VEGF) using immunomagnetic reduction (IMR). (A) Under external multiple ac magnetic fields, magnetic nanoparticles oscillate with the multiple ac magnetic fields via magnetic interaction. Thus, the reagent under external multiple ac magnetic fields shows a magnetic property, called mixed-frequency ac magnetic susceptibility. (B) Via the antibodies on the outmost shell, magnetic nanoparticles associate with magnetically label targets. With the association, magnetic nanoparticles become larger or clustered. The response of these larger magnetic nanoparticles to external multiple ac magnetic fields becomes much less than that of originally individual magnetic nanoparticles. (C) Detected VEGF concentrations in serum using SQUID-based IMR for normal people and patients with hepatocellular carcinoma (HCH) or colorectal cancer (CRC). (Reproduced with permission from Ref. [146], Copyright Hindawi Publishing Corporation, 2013.)

3.3 Atomic magnetometers (AM)

A rival magnetic detection technique to SQUID is atomic magnetometry. Atomic magnetometers (AM) are based on the measurement of the Larmor precession of spin-polarized atoms in a magnetic field [147]. The atoms are polarized by a polarized pump laser. Then, their precession in the magnetic field is detected by measuring the optical rotation or absorption of a probe laser, which may be the same laser used for pumping. Most AMs use the vapor of alkali metal (K, Rb, Cs) as resonant medium. In 1969, the first use of AM for detection of the static magnetic field produced by a sample of gaseous nuclear-polarized atoms was reported [148, 149]. Magnetometers using a 87Rb cell detected a 60-nG field produced by the nuclear spins with a sensitivity of 0.3 pT/Hz1/2. Owing to technological development of advanced tunable diode lasers and dense atomic vapor cell with a long relaxation time, modern atomic magnetometers (AM) have achieved as high sensitivity as most SQUID.

The sensitivity of AM depends on the product of three quantities in the following equation:

δB=gμBNτT

the magnetic moment of atoms (B), the number of atoms (N) involved in measurement, the spin-relaxation time (τ). As the collisions with the cell walls limit the spin-relaxation time, one of the most critical elements is paraffin-coated [150] or buffer-gas [151] alkali vapor cells to reduce wall relaxation. Currently, the most sensitive atomic magnetometer (AM) is the spin-exchange relaxation-free (SERF) magnetometer with the sensitivity of 0.5 fT/Hz1/2 [152]. The projected fundamental sensitivity limit of SERF magnetometer is on the order of 0.01 fT/Hz1/2, which is able to reach thermal magnetic noise of the human body (0.1 fT/Hz1/2) [116]. Therefore, with the advantage of ultrahigh sensitivity and not requiring cryogenic cooling, AM offer numerous potential applications in fundamental physics (symmetries of nature, electric dipole moment, charge-parity-time), geophysics (magnetic anomalies, dynamics of geomagnetic fields, magnetic properties of rocks), biomedical science (direct detection of biomagnetism and magnetic labels, detection of signals in NMR and MRI) [153].

For the biomedical applications, AMs have started to measure magnetic fields from the human heart in the 1970s [154]. A magnetocardiogram (MCG) system with AM was developed to detect the magnetic field from the heart and brain in a magnetically shielded environment [155]. A multi-channel SERF magnetometer has been used to simultaneously record MEG signals in a low ambient magnetic field environment [156]. A chip-scale AM was reported recently to detect MEG signals with a sensitive volume of 0.77 mm3 [157]. The first detection of ULF MRI with an atomic magnetometer (AM) was performed in 2006. By coupling with a remote-detection scheme, thereby improving the filling factor of the sample, time-resolved images of water flow were obtained with a temporal resolution of 0.1 s and spatial resolutions of 1.6 mm perpendicular to the flow and 4.5 mm along the flow [158]. To better understand the practical problems of applying AM in ULF MRI, MR image of a water phantom was studied by a portable atomic magnetometer working system with 2 mm image resolution [159].

There are two main methods for AM to detect magnetic particles. One is continuous flow carried by water. Estimated detection limits were 20 μm diameter for a single cobalt particle at a water flow rate of 30 ml/min, 5×103 magnetite particles (1 μm) at 160 ml/min, and 50 pl for the ferromagnetic fluid (10 nm) of cobalt nanoparticles at 130 ml/min [160]. Owing to the use of small vapor cells, an improvement by a factor of 105 in terms of the detected magnetic moment was reported with a microfabricated atomic magnetometer (AM) [161].

The second method is the scanning imaging scheme. From the full scanning profile, instead of a single-point measurement, the spatial information can be resolved, and the amount of the magnetic sample can be obtained simultaneously (Figure 10). The combination of approximately 20 μm resolution, nearly 1 cm detection distance, and picoliter-range detection limit makes this technique uniquely suited for practical applications of MNP in molecular imaging [162]. In addition, the AM used in this work had an operation temperature of 37°C, which is compatible with biomedical assay analysis and in vivo imaging [163]. With further development, two-dimensional scanning magnetic imaging of arbitrarily oriented MNP and quantitative molecular imaging of targeted antibody were demonstrated [164].

Scanning magnetic imaging of nanoparticles using atomic magnetometer. (A) Magnetic field profile of an assembly of approximately 7×105 magnetic nanoparticles (experimental data and fitted curve). The inset shows the geometry of the scanning axes with respect to the laser beam and the Cs detector. (B) A two-dimensional scanning magnetic image. Each stripe represents an x-axis scan. The width of each stripe is determined by the corresponding error bar of the fitted d value. (Reproduced with permission from Ref. [162], Copyright John Wiley and Sons, 2009.)
Figure 10

Scanning magnetic imaging of nanoparticles using atomic magnetometer. (A) Magnetic field profile of an assembly of approximately 7×105 magnetic nanoparticles (experimental data and fitted curve). The inset shows the geometry of the scanning axes with respect to the laser beam and the Cs detector. (B) A two-dimensional scanning magnetic image. Each stripe represents an x-axis scan. The width of each stripe is determined by the corresponding error bar of the fitted d value. (Reproduced with permission from Ref. [162], Copyright John Wiley and Sons, 2009.)

A recently developed technique based on AM solved the issue of lacking a spectroscopic parameter in magnetic sensing. This technique, termed as force-induced remnant magnetization spectroscopy (FIRMS), implements the binding force of the noncovalent bonds between the ligand on the MNPs and the targeted receptor molecules as a molecular signature for distinguishing different molecular interactions [165, 166]. The principle is shown in Figure 11. The FIRMS technique utilizes an external force with varying amplitudes to induce dissociation of the bonds between the target molecules on cells and the ligand molecules conjugated on magnetic labels. At each force, the magnetization and spatial coordinates of the magnetic particles with the particular binding character are revealed by scanning imaging of atomic magnetometer (AM). FIRMS has the ability not only to distinguish the molecule-specific binding from the nonspecific physisorption, but also to quantify the dissociation force of intermolecular noncovalent bonds (Figure 11), which is critical for the immunoassays, molecular recognition, and biomedical diagnostics. The well-defined binding force for the bonds between mouse IgG and magnetically labeled α-mouse IgG was calibrated to be 120 pN [167]. With the improvement of force resolution, two similar DNA double strands with one single-basepair difference were resolved [168].

FIRMS spectra and peak assignments. (A, B, C) FIRMS spectra showing the magnetization differential as a function of disturbing force for (A) blank experiment with no cells but only CD3 antibody and magnetic particles in the sample, (B) cell-binding experiment, where CD3+ T cells were incubated with CD3 antibody and magnetic particles, (C) control experiment, where CD3+ T cells that were not conjugated with the CD3 antibody were incubated with magnetic particles. (D, E, F) Schematics of binding behaviors corresponding to the peaks shown in the FIRMS spectra; (D) physisorption of the magnetic particles onto the sample well; (E) 70% specific binding to the target cells; (F) 45% non-specific binding of the magnetic particles on the cell surface. (Reproduced with permission from Ref. [165], Copyright John Wiley and Sons, 2011.)
Figure 11

FIRMS spectra and peak assignments. (A, B, C) FIRMS spectra showing the magnetization differential as a function of disturbing force for (A) blank experiment with no cells but only CD3 antibody and magnetic particles in the sample, (B) cell-binding experiment, where CD3+ T cells were incubated with CD3 antibody and magnetic particles, (C) control experiment, where CD3+ T cells that were not conjugated with the CD3 antibody were incubated with magnetic particles. (D, E, F) Schematics of binding behaviors corresponding to the peaks shown in the FIRMS spectra; (D) physisorption of the magnetic particles onto the sample well; (E) 70% specific binding to the target cells; (F) 45% non-specific binding of the magnetic particles on the cell surface. (Reproduced with permission from Ref. [165], Copyright John Wiley and Sons, 2011.)

A modified FIRMS technique was developed for microRNA detection in a label-free fashion. MicroRNAs are short, noncoding RNAs of approximately 20–25 nucleotides in length. Since their initial discovery two decades ago, microRNAs have been found to regulate about 30% of the human genome, involved in a lot of human diseases including cancer [169, 170]. The technique, exchange-induced remnant magnetization (EXIRM), also eliminates the requirement of using external forces. Instead, it uses the exchange reaction between the target microRNA and magnetically labeled RNA during competitive binding with the complementary sequence of the target microRNA. The combination of exchange reaction and quantitative molecular imaging with sensitive AMs enables detection of as few as 104 microRNA molecules [171].

FIRMS can also be used to probe the intrinsic mechanical property of biological molecules via labeling with MNP. In a recent publication, the binding forces of a series of DNA/RNA duplexes measured by FIRMS were used as internal force references to determine the intrinsic force generated by a motor protein, elongation factor G (EF-G) during the ribosome translocation, a key step in the process of protein synthesis. The results showed the EF-G power stroke to be 89 pN and it exerts for 0.5 nm [172].

3.4 Magnetic particle imaging (MPI)

Magnetic particle imaging (MPI) is a new medical imaging technique, which was invented in 2001 by Gleich and Weizenecker at Philips Research in Germany and was first reported in 2005 [173]. At present, MPI is in an exciting stage of commercial development like early MRI, when scanners and magnetic tracer (SPIO) have emerged as fast progressive development.

In contrast to both MRI, which measures the nuclear spin magnetization, and AM (and SQUID in many cases), which directly measures the remanent magnetic signal of the MNP, MPI takes advantage of the nonlinear response of SPIO to applied magnetic fields to obtain the magnetization of the sample, which can produce a tomographic image and quantify their local concentration. Under a strong magnetic field gradient, the SPIO is considered to be saturated at every point except the field free point (FFP), which defines the sensitive point. When a time-varying drive field of 100 Oe at 25.25 kHz is applied, the FFP scans rapidly across the field-of-view (FOV) of the sample. The SPIO located within the FFP respond to this rapid change in magnetic field by flipping their magnetization 180°. Because the receiver coil detect time-varying magnetizations, only the SPIOs at the FFP produce an MPI signal. Initial phantom experiments demonstrated a 2D spatial resolution of better than 1 mm and detection limit of 100 μmol/l for Fe, which is within the range of the allowed dosage for medical use [173].

Another main class of MPI scanner featuring two-dimensional projection mode was demonstrated recently [174]. Compared to FFP, the difference is that the strong selection field creates a zero field along a line, which has been termed as field free line (FFL). By rotating and translating the location of the projection FFL direction, an MPI dataset can be reconstructed to form a three-dimensional MPI image with nearly 10-fold SNR improvement [175].

MPI has extraordinary potential for molecular and cellular imaging because it has the advantages of high sensitivity and contrast, zero depth attenuation of signal, and accurate in vivo quantification. For example, shown in Figure 12, human embryonic stem cell (hESC)-derived cells were labeled with Resovist SPIO. When roughly twice the number of stem cells was injected into a postmortem mouse on the right site rather than the left, from the MPI image, the ratio of the signal intensities between the right and left injection sites was also found to be 2:1. The experimental detection limit is approximately 104 stem cells in the prototype system [176]. Similarly, encapsulation of SPIO into red blood cells (RBCs) has been suggested to increase the blood circulation time of nanoparticles. SPIO-loaded RBCs can be imaged using MPI in the blood pool of mice several hours after injection [177].

Stem cell imaging with MPI. (A) Projection MPI image of two injections of hESC-derived cells (1×105 cells on the left vs. 2×105 cells on the right site) introduced subdermally into a postmortem mouse (injection sites marked in cyan in the photo). The ratio of signal intensities between the right and left injection regions in the MPI image was found to be 2:1. Image acquisition time was 3 min total with FOV of 5×10 cm and 16 averages. (B) Plot of MPI signal intensity vs. number of stem cells in scanner, demonstrating excellent linear fit (R2=0.99). MPI images were acquired for nine stem cell populations ranging from 1×104 to 1×106 cells and compared for maximum signal intensity. Our current stem cell detection threshold (i.e., the noise floor) is limited by system interference and is approximately 1×104 cells in our prototype system. (Reproduced with permission from Ref. [176], Copyright Elsevier, 2013.)
Figure 12

Stem cell imaging with MPI. (A) Projection MPI image of two injections of hESC-derived cells (1×105 cells on the left vs. 2×105 cells on the right site) introduced subdermally into a postmortem mouse (injection sites marked in cyan in the photo). The ratio of signal intensities between the right and left injection regions in the MPI image was found to be 2:1. Image acquisition time was 3 min total with FOV of 5×10 cm and 16 averages. (B) Plot of MPI signal intensity vs. number of stem cells in scanner, demonstrating excellent linear fit (R2=0.99). MPI images were acquired for nine stem cell populations ranging from 1×104 to 1×106 cells and compared for maximum signal intensity. Our current stem cell detection threshold (i.e., the noise floor) is limited by system interference and is approximately 1×104 cells in our prototype system. (Reproduced with permission from Ref. [176], Copyright Elsevier, 2013.)

Moreover, because of its high temporal and spatial resolution, MPI has the potential application in clinical diagnostic imaging such as cardiovascular, neurologic, and peripheral vascular applications. At present, available magnetic tracers such as ferucarbotran can be administered intravenously. They are distributed in the vascular system for a limited time until the accumulation in the reticuloendothelial system. The first in vivo MPI studies have been presented by the Weizenecker group, which showed one axial slice of the anatomy and the cardiovascular system of a mouse [178]. Later dynamic results presented in vivo 3D real-time MPI scans exhibiting details of a beating mouse heart [179].

4 Conclusion

In this review, we have focused on magnetic detection techniques of MNP including HF (MRI) and ULF (GMR, SQUID, AM, MPI) and their applications in molecular imaging and diagnosis. MRI is a relatively mature technique that provides excellent spatial resolution of MNP-labeled molecules and cells. It also has a larger field-of-view compared to ULF techniques. New developments in MRI, for example, toward higher magnetic field, will certainly improve the spatial resolution and sensitivity as well. However, its sensitivity is usually lower than that of the ULF techniques; its bulky instrument and high cost also sometimes hinder its applications. In contrast, ULF techniques offer ultrahigh sensitivity, which makes them well suited for in vitro diagnosis. Their derived techniques, such as FIRMS and EXIRM, provides extra dimensions for molecular and cellular imaging, namely, molecular spectroscopy based on binding forces and label-free detection.

If we put these applications into the marketplace and establish the profitability of biomedical applications of MNPs to investors, the prospect for further scientific, technological, and commercial advances is indeed engaging.

Acknowledgments

This work was supported by the Institute of Chemistry, Chinese Academy of Sciences (Y41Z011BZ1).

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About the article

Li Yao

Li Yao is a Professor at the Institute of Chemistry, Chinese Academy of Sciences, where he received his doctorate in Physical Chemistry in 2008. His area of research includes nanobiotechnology, biophysical chemistry, and environmental biochemistry. His current research interests are focused on atomic magnetometry, high-performance magnetic prober, biomolecular recognition, and molecular and cellular magnetic imaging.

Shoujun Xu

Shoujun Xu is currently an Associate Professor of Chemistry at the University of Houston. He received his PhD in Physical Chemistry from Johns Hopkins University in 2002. His expertise includes spectroscopy, atomic magnetometry, nanobiotechnology, low-field magnetic resonance imaging, and magnetic sensing and imaging. Projects pursued in his laboratory include both technology development and fundamental biomedical research using the new techniques. Several techniques have been developed by him and colleagues, such as force-induced remnant magnetization spectroscopy. Fundamental research includes noncovalent bonding and mechanical force in biological processes.


Corresponding author: Li Yao, Beijing National Laboratory for Molecular Sciences, State Key Laboratory for Structural Chemistry of Unstable and Stable Species, Institute of Chemistry, Chinese Academy of Sciences, Beijing 100190, P.R. China, e-mail:


Received: 2013-12-25

Accepted: 2014-02-01

Published Online: 2014-03-22

Published in Print: 2014-06-01


Citation Information: Nanotechnology Reviews, Volume 3, Issue 3, Pages 247–268, ISSN (Online) 2191-9097, ISSN (Print) 2191-9089, DOI: https://doi.org/10.1515/ntrev-2013-0044.

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