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13th Asian Oceanian Congress of Radiology (AOCR), Taipei, Taiwan March 20-23, 2010

5th Congress of Asian Society of Cardiovascular Imaging, Hong Kong, 18-19 June 2011

Engineering and Physical Sciences in Medicine and the Australian Biomedical Engineering Conference, Australia, 14-18 August 2011

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Biomed Imaging Interv J 2010; 6(2):e13
doi: 10.2349/biij.6.2.e13
© 2010 Biomedical Imaging and Intervention Journal

PDF version Review Article

Development and use of iron oxide nanoparticles (Part 2): The application of iron oxide contrast agents in MRI

G Mandarano1, BAppSci (MedRad), GradCert (Higher Ed.), J Lodhia*,1, BAppSci (MedRadSci), P Eu2, BSci., MSci, NJ Ferris2, MBBS, MMed, FRANZCR, R Davidson1, PhD, MAppSci(MI), FIR, SF Cowell1, PhD, MEd, ANMT
1 Division of Medical Radiations, School of Medical Sciences, RMIT University, Victoria, Australia
2 Peter MacCallum Cancer Institute, Melbourne, Victoria, Australia


Magnetic resonance imaging (MRI) is a medical imaging tool that can incorporate contrast agents to enhance its ability to identify and characterise pathologies. MRI contrast agents can be paramagnetic such as gadolinium, or superparamagnetic such as iron oxide. Significant concerns of Nephrogenic Systemic Fibrosis (NSF) have arisen involving gadolinium-based contrast media.

Recent research has focused on iron oxide nanoparticles because their sizes are more comparable to biological units. These can give MRI the potential to detect a broader range of pathology, while also track and observe biological processes.

This is the second article of a two-part series and will review iron oxide nanoparticles as a MRI contrast agent, and the potential applications of iron oxide nanoparticles to a range of pathologies and processes involving MRI. � 2010 Biomedical Imaging and Intervention Journal. All rights reserved.

Keywords: Iron oxide nanoparticles, MRI


Since its clinical introduction, Magnetic Resonance Imaging (MRI) has been viewed as a highly advanced imaging modality. In particular, over the past decade, MRI has demonstrated its capability to generate images of anatomy and pathology with excellent contrast and spatial resolution. Also noteworthy has been MRI�s capability of imaging physiological processes with functional MRI (fMRI), Diffusion Weighted Imaging (DWI) and Perfusion Weighted Imaging (PWI). This of course has coincided with improved hardware and software developments. Overall, this has resulted in the medical community having a greater understanding and awareness of biological processes. Therefore, the clinical role and utility of MRI has evolved and is continually expanding.

Broadly defined, a contrast agent or medium is any substance that can be used together with an imaging technique to provide additional and useful information. Contrast media can be either exogenous or endogenous. Endogenous material which can be used include water molecules inherent within the blood stream when performing Arterial Spin Labelling (ASL) or Tagging (AST). Exogenous substances include the already well-known paramagnetic gadolinium-based agents and now emerging from literature, we are noting an increasing experimental usage of superparamagnetic iron oxide nanoparticles.

This is the second publication in a two-part series reviewing the potential use of iron oxide nanoparticles. Specifically, we will discuss the characteristics essential in iron oxide nanoparticles for MR imaging (as a contrast agent) as well as their potential clinical applications for a range of pathologies and physiological processes.

Contrast media in radiology and the current intravascular MRI contrast media

Intravascular contrast agents have been continually used since the early 1900�s across all imaging modalities in the field of radiology [1]. Over the past century, there has been a continual evolution of contrast media in medical imaging. These changes have been based on safety concerns (adverse reactions by patients), improved chemical technology (oil-based, water-soluble, ionic, non-ionic) and designing contrast agents dedicated to a particular imaging modality or technique [1, 2]. Additional concerns surrounding their implementation and use include the cost of contrast media, the need for certain contrast media preparations to be warmed to body temperature (viscosity, minimise adverse reactions) and also the need to have a recent reading of a patient's renal function (prior to administration of contrast media). There are also concerns surrounding compatibility, whereby a patient having an intravenous contrast media administered for a computed tomography (CT) study will be unable to have a nuclear medicine thyroid examination for several months; furthermore, this same contrast media is incompatible with metformin-based diabetic medication.

When MRI first became a clinical reality, it was thought that no contrast media would ever be needed because of MRI�s superior contrast and spatial (isotropic voxels) resolution compared with other modalities [3]. It was soon realised that a contrast media was needed to improve the specificity of MRI [4]. In 1988 the first MRI-specific contrast media preparation was approved by the Food and Drug Administration (FDA) for intravascular administration in clinical use [3, 5]. This was needed to define a pattern of contrast enhancement so that a characteristic enhancement pattern of a particular disease process could be recognised and also to narrow the differential diagnosis. This contrast media preparation contained gadolinium as its base. However, there have always been some concerns in relation to this preparation. These include its expense (cost per millilitre and patents are strongly held and continually renewed) and its degradation with exposure to ambient light.

The way that such gadolinium-based contrast media is chemically altered and eliminated by the body is not entirely understood [6]. For some time now, a condition known as �cross reactivity� has existed. It cannot be entirely explained, however, it is thought to result either from the chelated molecules or elements, or the chelated structure themselves [7].

Within the last several years, a more serious condition has been attributed to gadolinium-based preparations. This condition is referred to as nephrogenic systemic fibrosis (NSF) and can lead to eventual death. This condition is almost always seen in patients with reduced renal function (less than normal glomeruli filtration rate, GFR) and there have been a number of deaths recorded and attributed to NSF. In early 2008, the Royal Australian and New Zealand College of Radiologists (RANZCR) has recommended that all clinical centres offering gadolinium-based contrast media for MRI scanning examinations to establish a new policy concerning intravascular administration and, in particular, with respect to NSF and patients with impaired renal function [8]. Due to the above trends, anecdotal reports suggest more caution and less reliance upon gadolinium-based contrast media even though no alternative currently exists in Australia.

The current commercially available contrast media is gadolinium-based and also referred to as para-magnetic and only benefits T1-weighted MR imaging. No mainstream commercially available contrast media for T2-weighted imaging is currently available in Australia. There are benefits attributed to contrast media if it can be prepared for T2-weighted MR imaging [6]. Coincidently, an increasing trend is underway towards clinical MRI scanners with higher field strengths such as 3 Tesla (6). The main drivers are improved capital and running costs and increased signal-to-noise (SNR) ratio. However, at higher field strengths such as 3 Tesla, T1 and T2 relaxation times of human tissue are altered compared to 1.5 Tesla. Gadolinium-based contrast media �works� by shortening T1 tissue relaxivity values and therefore only T1 optimised sequences can be used (T2 relaxivity values are not altered significantly for MR imaging practicality).

It is proposed that contrast media preparations based on nanoparticles can overcome all of the abovementioned challenges related to MRI scanning while simultaneously addressing the current medical safety concerns [6, 9]. More specifically, iron oxide-based nanoparticles have the following attractive attributes:

  • it can offer T2-weighted imaging opportunities
  • it has a well-recognised and understood pathway for breakdown and excretion from the human body
  •         Degradation causes iron to enter plasma, where it is processed by the body

            Risk of iron overload is minimal

            Average dose of iron in contrast agent is comparable to iron contained in less than one unit of blood

  • It provides �negative� enhancement

Depending on the physio-chemical property of the coating (surrounding the iron oxide nanoparticle), both generalised and specific contrast media can be created [9, 10]. The term �specific� means that contrast media preparations can be targeted to a particular organ within the body or a particular disease process. If this can be achieved, then it follows that not only can diagnostic imaging be successful, but also therapeutic drugs/medication can be tagged to the preparation so that it can reach and work on the target tissue.

At one point, it was estimated that 30% to 40% of MR examinations were performed with intravascular contrast media [7]. With the awareness of NSF, no hard data currently exists to determine if the use of intravascular contrast media has decreased or remains at the same level.

The most commonly available intravenous contrast media contains gadolinium. A gadolinium ion has seven unpaired electrons in its outer shell [11] and is considered a paramagnetic substance because it has an overall positive effect on the local magnetic field [3]. In brief, when placed within a magnetic field, the negatively charged gadolinium ion demonstrates characteristics such as a magnetic moment, producing a large time-varying magnetic field in its vicinity, allowing rapid exchange of bulk water, altering the relaxation rates (both T1 and T2, or longitudinal and horizontal) of adjacent water protons [7, 11, 12]. Gadolinium is referred to as a T1 enhancement contrast agent as it affects T1 to a greater extent than it does T2. The act of molecular tumbling and local magnetic field alterations occur near the Larmor frequency value. This leads to a reduction of the T1 relaxation value of adjacent water protons, which in turn, leads to an increased signal strength on T1 weighted images [11]. This is due to an increased rate of longitudinal magnetisation recovery [7]. For acceptable biocompatibility, gadolinium is chelated to other molecules. This reduces any acute toxicity effects, and also allows the gadolinium-based agent to remain circulating within the body for a relatively longer period (than without chelation) [12] with an elimination half life of 1 hour to 2 hours.

Currently available paramagnetic contrast agents are commonly administered intravenously. Its biodistribution is into the blood stream and then into the extra cellular space. It is therefore not taken up by any specific body organ, tissue type or pathologic lesion. Hence, gadolinium compounds are also regarded as non-specific contrast agents [12]. However, enhancement patterns are known to be characteristic of certain pathology groups. For example, a hyperintense circular rim with a hypointense centre may be representative of a cystic lesion.

Current concerns with nephrogenic systemic fibrosis and gadolinium

Up until about ten years ago, gadolinium-based contrast agents have been regarded as having a relatively excellent safety record [13]. NSF was originally referred to as nephrogenic fibrosing dermopathy by Cowper et al. in 2000 [14]. It was described as being scleromyxoedema-like cutaneous disorder and thought to only affect the skin or dermis. It was noted in patients undergoing renal dialysis [14, 15]. As additional cases became recorded and further understanding of the pathology grew, the currently used term of NSF has become accepted. This is due to the now recognised systemic nature of the pathology [16, 17]. Commonly, NSF commences with swelling at the distal aspects of the extremeties. This may then resolve, however, leaving behind thick, firm plaques over the affected skin.

In the majority of patients, initial skin lesions appear on the legs, then the arms and lastly on the trunk of the body. It has also been reported that the skin lesions are often symmetrical and bilateral. It may then progress to a point where the patient has significantly reduced range of motion of their extremities and joints [18]. In addition to the flexion and joint contractures accompanying extremity skin lesions [19, 20], fibrotic effects may also be widespread and penetrating; involving organs including the liver, lungs and heart, among others [20, 21].

Today, the only successful approach in treating NSF is to restore the normal renal function. This can only be achieved by renal transplant surgery [18, 20]. NSF is almost always seen in patients with less than normal renal function or patients requiring ongoing renal dialysis. Therefore, NSF may be a resulting consequence in patients with renal impairment because the contrast media excretion half life is markedly increased [22]. This situation then permits disassociated or de-chelated gadolinium ions an increased circulation time. Some authors consider NSF to be an, �adverse reaction to gadolinium contrast agents in particular the less stable gadodiamide� [23].

Even though all gadolinium-based preparations carry a level of risk, there is published evidence to suggest that some gadolimium-based contrast agents offer a greater risk than others in inducing NSF [23]. This variation in risk is linked to the overall molecular structure of the gadolinium-based contrast agent and in particular, its level of chemical stability within the human body [24]. Gadodiamide has been associated with the greatest incidence of induced NSF [23, 24]. On its own, gadolinium is highly toxic. Not only can it cause injury to the liver and spleen, but it can also inhibit secretions of certain enzymes and it can induce haematological ailments [24-26]. To minimise such toxic consequences, the gadolinium ion is chelated to other chemical elements and compounds [24]. This improves its biocompatibility. The molecule or atom that is bonded to the gadolinium can be referred to as a ligand. The gadolinium ion is chelated in either a linear or macrocyclic fashion and prepared as either ionic or non-ionic formulations [7, 24]. Published data reflect that the least stable preparations are non-ionic linear chelate formulations such as gadodiamide (Omniscan, GE healthcare, Chalfort, ST Giles, UK) and Gadoversetamide (OptiMark, Covidien, St Louis, USA). Gadodiamide has been reported to have a kinetic stability of 35 seconds. That is, at a pH of 1.0, half of the preparation will shed the linearly chelated material; thus leaving free gadolinium ions to search for other metals (body cations) to bind to within the body. These would include iron, copper, zinc and calcium. This process is referred to as transmetallation. Of the above body cations, zinc has the highest relative concentration (55-125 micromole per litre) within the blood stream [24].

The most stable gadolinium-based preparation is the ionic, macrocyclic chelate formulation; namely, gadoterate. As of April 2009, no cases of NSF linked to any macrocyclic formulation has been reported [24] or confirmed by the International Centre For Nephrogenic Fibrosing Dermopathy Research [27].

This is because a macrocyclic structure provides relatively superior protection of the gadolinium ion. That is, the gadolinium ion is caged by the chelating agent [7, 24, 28, 29]. Conversely, a linear chelate is referred to as being a flexible open chain and thus not providing a strong bond to the gadolinium ion. Gadoterate is documented to have kinetic stability of greater than one month [24].

High et al. [13] obtained paraffin embedded tissue samples from the NSF registry (the International Centre for Nephrogenic Fibrosing Dermopathy Research). These tissue samples had histopathologic diagnosis of NSF. This research group demonstrated with energy dispersive spectroscopy (EDS), a device used to characterise chemical elements, that in four of seven patients, gadolinium was able to be identified and all detectable gadolinium particles were less than 1 micrometre in size. Further analysis with field emission scanning electron microscope (FESEM) demonstrated that in all of the positive tissue samples, gadolinium particles were present within the intracellular space and most probably located within, or adjacent to, the lysosome structures. Also noteworthy was an excessive amount of iron deposition within the tissue samples.

While the exact cause of NSF has not been conclusively established [20, 21, 29-33] and precise pathologic pathways are yet to be determined [15], there is however, convincing evidence that gadolinium may be responsible somehow [32, 34]. The most probable theory is that de-chelation occurs [35], resulting in the release of free gadolinium ions, which in turn may or may not lead to transmetallation [36, 37].

Iron oxide nanoparticles as MRI contrast agents

The most common form of iron oxides used in nanoparticle preparations are magnetite (Fe3O4) and maghemite (γFe2O3) [38, 39], and research with these has been intensive for about a decade now. They are both insoluble in water and because of their size, these superparamagnetic substances only exhibit their magnetic properties when placed within a magnetic field [12, 40]. The hydrodynamic size of a nanoparticle preparation is the term used to describe its overall size, that is, the iron oxide core plus the coating plus any additional ligand attachments (see Figure 1). If the overall hydrodynamic size is greater than 50nm, then the preparation is referred to as a superparamagnetic iron oxide nanoparticle (SPION). If the hydrodynamic size is less than this, then the preparation is termed ultra small superparamagnetic iron oxide (USPION). For the purpose of this manuscript, the authors will use the term iron oxide nanoparticle (IONP) as a generic term to refer to nanosystems containing a core of iron oxide.

Iron oxide nanoparticle preparations are highly complex. There are numerous production methods which can be used to generate them. Each method can result in iron oxide nanoparticles having specific dimensions, as well as unique imaging and therapeutic characteristics. Each manufacturing method is undertaken in a strict controlled environment to ensure consistency of dimensions, characteristics and biostability.

The physiochemical properties of iron oxide nanoparticles are determined by the size of the iron oxide core, its overall charge and the zeta charges between coatings and the overall hydrodynamic size. With respect to magnetic resonance imaging, the above factors also play a fundamental role in determining their efficacy (or imaging efficacy), stability within the body�s environment, biodistribution, opsonisation, metabolism, clearance from vascular system and then excretion from the body [41].

Part one of this journal article series discussed the variety of production methods. The advantages and disadvantages of this method were also presented and therefore that information will not be presented here.

The coating

An understanding of the bonding and the geometry of the coating will help us appreciate the pharmokinetic pathways and biodistribution of contrast agents composed with iron oxide nanoparticles [40, 42-44]. From a chemical perspective, the iron oxide core is coated for four main reasons. Firstly, to prevent destabilisation; secondly, to prevent agglomeration (aggregation or sedimentation) as it will be a colloidal suspension; thirdly, it allows for the iron oxide nanoparticle formulation to be soluble in an aqueous solution or a biological medium; fourthly, it determines either the role it performs within the body (diagnostic magnetic resonance imaging, cell tracking or therapeutic purposes such as tailored drug delivery) or the ligand that can be bonded to it to support the imaging, tracking or therapeutic roles.

The coating used can also facilitate the method of endocytosis [45]. For example, it has been shown that IONP coated with monomer citrate (overall hydrodynamic size of 8nm) demonstrated cell entry via phagocytosis. When the same iron oxide nanoparticles were coated with polymer carboxydextran (overall hydrodynabic size of 31nm), cell entry or penetration was demonstrated by pinocytosis. In both examples, the same cell line was used.


In the literature, high density coatings have been reported to be effective in stabilising iron oxide nanoparticles [40, 46-48]. Such high density coatings are commonly polymeric and monomeric materials or species. Polymeric coatings include dextran, carboxymethylated dextran, carboxy dextran and starch. Whereas monomeric coatings include dimercaptosuccinic acid (DMSA), amino acids and α‑hydroxamates (such as citric, tartaric or gluconic acids).

The Hydrodynamic Size

There is evidence to suggest that USPION is less prone to phagocytosis by the liver; whereas SPION greater than 50nm are rapidly phagocytosed [40]. Therefore, hydrodynamic size can affect biodistribution and blood half life in a time dependent manner [49]. Eventually, USPIO will actually be processed by the liver.


Iron oxide nanoparticles greater than 50nm are readily macrophaged by the reticuloendothelial system (RES). This namely refers to the Kuppfer cells of the liver, the spleen and bone marrow [40]. Iron oxide nanoparticles less than 50nm have been used to demonstrate uptake by lymph nodes [40, 41, 50-53]. Of the available iron oxide-based contrast agents which are currently on the market and also undergoing clinical trials, the blood half life values can vary considerably from 40 mins to up to 36 hours. There is a link between the hydrodynamic size and the biodistribution and blood half life.

As has been established, particles greater than 50nm are readily taken up by the liver in a matter of minutes. USPION are not readily phagocytosed by the liver and can have a longer blood half life and thus reach other structures within the body [40]. It must also be emphasised that the coating itself can be responsible for aspects of biodistribution as well as ligands; for example the targeting of specific cells or organs [40, 41, 51].

Metabolism and excretion

The manner in which the body metabolises iron oxide nanoparticles is determined by their overall chemical composition. In particular the immediate coating and any ligands are strong determinants as to the site (that is, which particular organ or body system) of metabolism and thus also the rate of metabolism and excretory pathways.

The commonly used dextran coating should ideally be of low molecular weight. This is vital as higher molecular weight dextrans, such as those used as plasma substitutes, have a reported association with adverse reactions. Immunoglobulin G (IgG) antibodies have been reported to be reactive to such high molecular weight dextrans [54]. Low molecular weight dextrans will initially undergo dextranases which is an intracellular level breakdown process. The majority of the breakdown components are excreted with urine over a period of nearly 2 months [40]. Once the low molecular dextran has metabolised in this manner, the iron oxide core has been found to enter the normal iron store of the body. Such iron elements can also be found as haemoglobin with the body�s red blood cells [55]. This iron then follows the same excretory pathway as endogenous iron. That is, approximately one-fifth is eliminated mostly with the faeces and over a three-month period. Therefore, as both dextran and iron from iron oxide nanoparticles are incorporated into the body�s normal metabolic pathways, without raising these levels noticeably, and with the evidence available today, it can thus be stated that these substances do not trigger any long-term toxicity.

In the average healthy adult, normal total human iron stores is about 3500mg with the liver containing an average of 0.2mg of iron per gram [55]. From the currently approved iron oxide nanoparticles for diagnostic MR imaging, a regular adult dose can contain 50-200mg of iron. This value can be considered relatively small compared to the human body�s iron store. Chronic iron toxicity is known to occur when the concentration of iron within the liver reaches a level of 4mg of iron per gram of liver [38].

MRI imaging with iron oxide nanoparticles

Phenomenon of superparamagnetism

Iron oxide nanoparticles, as discussed here, are referred to as being superparamagnetic. The superparamagnetic phenomenon is observed when the thermal energy of the medium is sufficient to alter the crystallite or nanoparticles�s magnetisation direction by overcoming coupling forces. When crystallite or nanoparticles are placed within an external magnetic field, its magnetic moment will align with the externally applied magnetic field.

At least two points distinguish superparamagnetism from paramagnetism. Firstly, with paramagnetism, it is each individual atom or ion that becomes aligned with an externally applied magnetic field. Secondly, superparamagnetism will occur when the crystal or hydrodynamic size is less than its ferromagnetic domain. Authors report this size to be less than 30nm [40], with the �critical size� being about 15nm [44, 56, 57].

When not in the presence of an externally applied magnetic field, superparamagnetic particles are not magnetised, nor do they demonstrate any remnents of magnetism once removed from the magnetic field. When the crystals or particles are under the influence of an applied magnetic field, their magnetic spins are considered to be in perfect alignment and very high local magnetic field gradients are generated. These gradients then cause spin dephasing of the surrounding water protons; thus reducing their T1 and T2 relaxivity [12, 56, 58].

It must also be noted that there is a relationship between the iron oxide nanoparticle size and the level of superparamagnetic saturation. As the particle size decreases, so too does the superparamagnetic saturation. This then has an effect on reducing the observed relaxivity or further reducing T1 and T2 relaxation.

IONP Influence on magnetic resonance image characteristics

Compared to a paramagnetic material such as gadolinium, the relatively high magnetic moment of superparamagnetic species, such as iron oxides, are sometimes referred to as super spins [40, 59]. The dipolar interactions between the super spins and adjacent water protons result in both high longitudinal (r1) and transverse (r2) relaxation values. IONPs therefore increase T2* relaxation rates through the susceptibility effect and thus have their greatest visual impact on T2*‑weighted images produced with gradient echo-based pulsed sequences [48, 60-62]. The accelerated phase loss due to local field gradients generated by super spins, all stem from the (induced magnetisation) high susceptibility level of iron oxide.

At the common clinically available field strengths of 1.5T and 3.0T, published data [40] indicate that any aggregation of SPION will only slightly decrease r1 (longitudinal relaxation) and dramatically increase r2 (horizontal relaxation).

Magnetic resonance imaging with iron oxide nanoparticles

Even though images composed with gradient echo pulse sequences posses lower signal-to-noise ratio and spatial resolution compared to spin echo pulse sequences, they are currently the most appropriate imaging sequence to use with SPIONs. This is often termed magnetic susceptibility imaging [4, 12, 48, 63]. The contrast enhancement captured on an MR image is dependant upon a number of factors. Most notable are the biodistribution and opsonisation of SPIONs.

Iron oxide nanoparticles as contrast agents for MRI

In addition to their superior biocompatibility, IONP MRI contrast agents have been documented to increase diagnostic sensitivity and specificity [41, 45, 64-66] in both animal model experimentation and in human trials.

This improved accuracy has been attributed to their superparamagnetic effects and relaxation times [41, 64, 66]. Efficacy of IONPs as MRI dedicated contrast media also largely depends upon their physiochemical properties [41]. Such attributes include: size (both of the iron oxide core and the overall hydrodynamic dimensions); coating (dextran derivative or other); and the zeta surface charges. Their efficacy can be further increased with complex surface modifications. This is achieved by bonding or attaching active material such as monoclonal antibodies, receptor ligands and also proteins [41].

For intravascular administration, the hydrodynamic diameter of IONPs are very rarely greater than 150nm. The iron oxide core itself is usually no more than about 15nm [67]. The coating itself is preferably composed of dextran, or a derivative, of a low molecular weight. These are positive properties, as the dextran is biodegradable and its low molecular weight minimises possibilities of adverse reactions [54]. IONPs have also been incorporated into oral contrast media [41].

Imaging challenges to consider and overcome

There are a few imaging challenges with the use of IONPs [68]. They are in relation to commonly encountered artefacts in MRI imaging. On their own, they can be frustrating to deal with in everyday imaging. However, when IONPs are included in the imaging regime, a further layer of complication is added.

The first criticism is that it is difficult to determine or differentiate a signal void induced by IONPs compared to signal voids generated as artefacts by materials such as metal (susceptibility artefact). The second artefact is that of partial volume averaging. IONPs are capable of being involved in processes occurring at a cellular level. Hence, signal voids induced by IONPs that are smaller than the spatial resolution of the MRI image will not be represented as distinct and within a voxel; as individual signal intensities within a voxel are averaged together [4, 11, 12, 68].

Emerging trends: applications of iron oxide nanoparticles and the role of MRI

Where possible and practicable, medical investigators and clinicians would prefer an investigation means that is non-invasive or minimally invasive. This approach is safer for patients, it expedites the medical management of patients, and can negate morbidity and mortality consequences. IONP preparations, combined with MRI, have the potential to revolutionise a number of investigative and treatment procedures. This would be achieved by combining the advantages provided by IONP together with MRI, leading to safer and superior imaging alternatives. IONPs as MRI contrast agents have already been discussed in this manuscript. To re-iterate, they promise improved levels of toxicity and increased diagnostic sensitivity and specificity. These are achieved through careful chemical preparations, leading to IONPs having the required characteristics for biocompatibility and MRI image enhancement. It is recognised that further research is required to overcome the challenges mentioned.

We now follow with a discussion on the innovative uses of IONPs combined with MR techniques. These promise to revolutionise clinical therapies and improve patient outcomes.

Molecular imaging is a broad term concerning the imaging of biological events at the cellular or molecular level. It should also be non-invasive and the imaging characteristics representing the biological activity should be quantifiable [69]. MRI is seen as having an emerging and innovative role. Molecular imaging can be possible with MRI when IONPs are conjugated with biologically active materials such as antibodies.

The near future looks promising for MRI, together with IONPs, to have a positive impact in leading non-invasive imaging of biological and biochemical processes. Not only can such processes be diagnosed; but also progression and treatment can be imaged over time.


Angiogenesis, the growth of new blood vessels (for development, wound repair or tissue reproduction), is related to tumour growth and progression [70]. There are several known molecular markers associated with angiogenesis. That is, endothelial cells active in angiogenesis express known surface receptors compared to endothelial cells not partaking in angiogenesis [71]. The commonly occurring receptors include integrins and vascular endothelial growth factor receptors. Antibodies or drugs to seek out angiogenic markers, can be conjugated to IONPs and imaged with MRI. Thus, the angiogenic process can be identified and any success in treatment can be accurately monitored. This can be achieved by exploiting the increased permeability of newly formed tumour vessels compared to normal healthy vessels [72].


Apoptosis is the self destruction of cells. When determined by cell age or cell health status, the nucleus triggers this process. It requires metabolic activity by the dying cell and is commonly characterised by a redistribution of phosphatidylsenine in the cell membrane [70]. It can even be associated with tissue development and homeostasis [71].

The degree of apoptosis can determine how successful chemotherapy and radiation therapy can be. Identifying apoptotic events in vivo would hence further evaluate treatment regimes and progression of pathology [71]. It is known that apoptotic cells express lipid phospatidylserine (a phospholipid) on their cell membrane. Synaptotagmin I is a protein that is used to detect this phospholipid. When this protein was conjugated to IONPs, apoptosis was demonstrated in vivo with mice [73].

The capability to image apoptosis can allow for almost real-time monitoring of efficacy of drug therapies [56].

Targeted drug delivery with IONP and MRI

Many therapeutic drugs that are available, can be considered non-specific in nature. By non-specific, it is meant that such drugs are administered intravascularly and are thus distributed randomly. This can lead to unwanted effects on healthy tissue [74]. Specificity for target tissue or cells can be achieved by conjugating IONPs with ligands [71]. Such ligands include antibodies (in particular for targeting cancerous tissue or cells), proteins, peptides and other biological markers.

Targeted drug delivery, as provided by superparamagnetic colloid suspensions, can be guided by an external magnetic field to the site of interest [41], thus, having the capability to minimise both side effects and required dose [57, 74, 75]. Therefore, pharmaceutical drugs can be binded to IONPs designed to reach a specific, or target, organ, and then be released there [41, 44]. The emerging breakthroughs that make magnetic drug targeting possible and promising are the new classes of IONPs less than 50nm. This allows for improved circulation time, thereby permitting delivery to the target site without the likelihood of being sequestrated by the RES before this can occur [41]. With the original classes of IONPs that became commercially available over ten years ago, RES uptake occurred within a few minutes following intravascular administration (hence, their original application as dedicated contrast agents for the liver) [41].

Drug targeting can be achieved by passive, active or physical means [41]. Magnetic drug targeting falls into the category of physical means; as the pharmaceutical is attached to a carrier system (the IONP) and its distribution is facilitated with an external influence (the magnetic field).

Not all therapeutic drugs can be conjugated to one single variety of IONP. Characteristics of IONPs that can determine attachment of therapeutic drugs include their size, surface charges (zeta charges) and capacity for protein absorption [41]. The process of cell uptake is determined by the overall size of the nano-system (IONP, surface coating/s and pharmaceutical); phagocytosis or pinocytosis. Pinocytosis occurs for items less than 150nm [41, 74]. The condition under which a cell finds itself in, may alter its susceptibility to a nano-system. For example, under normal conditions, walls of endothelial cells are permeable to objects 10nm or smaller. However, when involved in pathologic processes such as tumour infiltration and inflammation, the endothelial wall can be permeable to objects up to as large as 700nm [41]. Zeta charges need to be carefully managed. They determine whether or not nanoparticles aggregate or if they remain suspended in its medium. More importantly, they also play a part in endocytosis. It is noted that the likelihood of phagocytosis increases with a higher zeta charge [41, 76], while time spent within the circulatory system is reduced. There is an electrostatic process involved when particles and substances are absorbed by a cell�s outer membrane [77]. Understanding nanosystem interaction with proteins is vital, as when they are injected into the vascular system, their first interaction is with the plasma proteins. Therefore, the manner in which nanosystems are capable of interacting with opsonins (proteins that encourage phagocytosis such as IgG) and dysopsonins (proteins inhibiting phagocytosis) also determine if they are readily phagocytosed by the RES or if they can reach their intended target and release their pharmaceuticals. Hence, protein repulsive molecules such as polyethylene glycol (PEG) can be used to modify the surface of nano-systems to reduce their recognition by the RES [74] and reduce non-specific cellular uptake [78].

A phase I/II clinical trial of IONPs combined with epirubicin designed to image and treat solid tumours (such as sarcomas), showed that these nanosystems were reasonably well tolerated by the fourteen patients involved. No organ toxicity attributable to iron oxide was noted. However, toxicity responses to epirubicin were recorded at doses greater than 50 mg/m3 [79].

Thermal applications for cancer cells: magnetocytolysis and hyperthermia with IONP and MRI

Compared to normal healthy cells, cancer cells are known to be sensitive to temperatures above 42 degrees Celsius. Normal cells can survive at higher temperatures [44]. In cancer cells, at temperatures above 42 degrees Celsius, protein function is disrupted which can lead to apoptosis [80]. Thus, hyperthermia is a proposed treatment regime for certain cancers. Until recently, hyperthermic approaches have included irradiation with radiofrequency, ultrasound and microwaves. One known criticism of these approaches is the likelihood of hyperthermic injury extending to healthy tissue. The term, magnetic induction hyperthermia, now refers to cancer tissue being exposed to an alternating magnetic field [44].

Hyperthermia using IONPs together with MRI has demonstrated positive results in pre-clinical evaluation studies. With this combined approach, magnetic nanoparticles can be either directly injected into a tumour volume or designed to be selectively uptaken by a tumour site. This target-selective capability improves local heating treatment to the tumour while dramatically minimising potential for injury to surrounding healthy tissue [81]. Furthermore, the alternating magnetic field (not absorbed by tissue), together with appropriately prepared IONP, can allow hyperthermia treatment [41] to be applied to areas deep within the body [44]. Radiofrequency pulses provided by MRI can be designed to provide frequencies and amplitudes to increase local cell temperature up to 55 degrees Celsius [80], thereby inducing cytolysis. This process, therefore, can be used to generate heat to target cells [82].

However, for magnetic hyperthermia to be successful, it requires accurate delivery of magnetic nanoparticles to the tumour site.

A number of experiments report successful use of magnetic induction hyperthermia in cancerous cell models and also in animal models [44, 81, 83].

Salado et al. [84] successfully demonstrated in a rat model, with in vivo MRI imaging, that the IONPs which they developed were capable of providing positive contrast enhancement of induced liver tumour and also successfully treated these liver tumours with MRI-induced hyperthermia. Thereby, demonstrating that IONPs can have both a diagnostic and therapeutic use. Analysis of the rats following the experimental study demonstrated no vascular embolisms (the IONP preparation was injected through the ileo-colic vein) and specimens of the liver demonstrated insignificant inflammatory changes.

Xu et al. [85] have produced nanoparticles containing a core composed of iron and cobalt and a gold shell. These nanoparticles demonstrated a specific magnetisation value far greater (226 emu/g) than that achievable with commercially available iron oxides (78.8 emu/g). This higher magnetic moment is claimed to improve heating efficiency in hyperthermia applications. However, their publication did not mention any results or discussion of toxicity or biocompatibility studies.

Initial success of magnetic hyperthermia with small groups of human patients provides a promising outlook for future clinical applications. Plotkin et al. [86] reports a study on eleven consecutive patients (mean age 44 years), each with recurrent supratentorial glioblastoma. All patients had previously undergone surgery, nine patients also had radiation therapy and eight patients also had chemotherapy. Based on the prognosis following these treatments, these eleven patients were eligible as candidates for hyperthermia using nanoparticles and MRI. Nanoparticles were then administered directly into the tumour volume. MRI hyperthermia, or nano cancer therapy, followed and in ten patients, the mean reduction in gross tumour volume (GTV) was 74% as indicated with PET-CT fusion imaging.

Cancer imaging with IONP and MRI

Diagnosing cancer in its early stages significantly improves patient outcomes and survival rates. The initial use for IONPs was directed at imaging liver tumours [87]. This was due to the nanoparticles being greater than 60nm and therefore readily phagocytosed by the liver. This has been occurring for several years now and there are a few commercially available preparations for this specific purpose. The authors will therefore discuss IONPs in relation to other cancers.

Current MRI techniques allow for the detection of tumour sizes in the order of one centimetre cubed. By conjugating known cancer antibodies with IONPs, then MRI can be used to identify cancerous tissue of smaller dimensions through molecular interactions. This is an improved sensitivity for cancer markers, compared with current cancer marker detection probes [88].

IONP can be coated with (DMSA), a bi-functional chelating agent and ligand. Herceptin, a monoclonal antibody, uses elements from within the immune system to stop tumour progression by binding to HER2 receptor and triggering a response by natural killer (NK) cells [89]. With a 1.5 Tesla clinical MRI scanner, Lee et al. [88] successfully demonstrated how to identify cancer sizes as small as 50mg in mice using IONPs conjugated with Herceptin.

The above approach improves patient outcomes compared with just chemotherapy alone.

This principle has also been used to target other tumour antigens. IONPs conjugated with peptide EPPT1 (synthetic peptide EPPT1, also known as alpha-M2 peptide (YCAREPPTRTFAYWG), derived from the CDR3 Vh region of a monoclonal antibody, ASM2) are able to target underglycosylated MUC-1 (mucin 1); which is a tumour antigen expressed by many epithelial cell adenocarcinomas such as pancreatic, colorectal, gastric and prostate [56].

The role that lymph nodes play in cancer staging has not gone unnoticed by researchers in this field. IONPs and MRI can be used to image the condition of lymph nodes and more accurately determine the extent of metastatic spread [69]. Oghabian et al. [90] demonstrated 98% detection sensitivity with in vivo imaging of a rat model at 1.5 Tesla. They also concluded that the type of surface coating and its thickness were factors in determining MRI signal intensity.

Today, prostate cancer is still a leading cause of death in men. Current treatment, among others, involves brachytherapy, and as a procedure, it has its own level of invasiveness, morbidity and mortality. Wang et al. [91] have conjugated IONPs with prostate specific membrane antigen (PSMA) and also with doxorubicin, a chemotherapy drug. By performing whole cell assays on human cell lines, Wang et al. demonstrated that their conjugate can detect, with high sensitivity, prostate cancer cells expressing PSMA, thereby, promising the possibility of a multifunctional (diagnostic and therapeutic) nanoparticle system.

IONPs can also be used to better identify tumour boundaries within the brain [69, 87, 92, 93]. This leads to improved quantification of tumour volumes. Compared with gadolinium-based contrast agents for MRI and also taking into consideration oedema surrounding a tumour volume, IONPs can define tumour margins for longer time periods [69, 92, 94]. This is as a result of the IONPs being endocytosed by tumour cells. Thus being internalised, their elimination rate from the tumour is longer compared to extracellular gadolinium-based contrast media.

Cell labelling and tracking; including stem cell therapies

MRI imaging of cells labelled or endocytosed with IONPs is considered an indirect imaging technique. This is because changes in MRI signal intensity is in relation to the amount of IONPs and not due to the number of cells. One concern is the MRI signal characteristic and change over a time period. As stem cells rapidly divide, the fixed amount of IONPs is spread throughout the volume of newly divided daughter cells [68, 75]. This will be a relative decrease in MRI signal not accurately reflecting the activity of cells. The second noteworthy concern is the possibility of false positive signal findings. This is a result of iron presented from cells undergoing apoptosis or lysis [68]. Despite these challenges, possibilities are being created for many biomedical applications [77].

The possibility of imaging stem cells in therapeutic applications with MRI is becoming an ever increasing area of active research. Published research so far indicates that IONPs combined with peptides or transfection agents can be used for stem cell uptake [75].

A study using stem cells with IONPs injected into the infarcted myocardium of pigs was able to be successfully imaged at 1.5 Tesla in vivo [95]. Following this, histological analysis revealed that the MRI signal appearance attributed to IONPs corresponded with stem cells that had taken up IONPs.

Heymer et al. [96] combined human mesenchymal stem cells (hMSCs) with IONPs; placed these in collage type I hydrogel (clinically approved for the repair of articular cartilage) and imaged them in a 11.7 Tesla MRI scanner. Iron uptake was confirmed by histological analysis and correlated with hypo-intense regions demonstrated on the MRI images. Their technique offers the possibility to use MRI to track the migration of IONPs loaded with hMSCs following implantation for articular cartilage repair.

Stem cell research is highly regarded as offering possible treatment solutions for patients with neuronal pathologies and injuries. Guzman et al. [97] proved that IONPs themselves do not alter survival rates, migration patterns or differentiation capabilities of stem cells from the human central nervous system (CNS), compared with unlabelled human CNS stem cells. They demonstrated this by administering the combined human CNS stem cells and IONPs into neonatal, adult and injured rodent brains. They used MRI to track the migration of stem cells and confirmed the image findings histologically.

Stem cells labelled with IONPs have been widely used in animal models (mice, rats and pigs) to demonstrate regeneration of the myocardium following infarction [98]. The limitation of this technique is that the conjugation of IONPs and stem cells need to be injected directly into the myocardium. This introduces an element of invasiveness. However, the region of infarct can be clearly delineated [99]. So far, administration of stem cell and IONP conjugations for myocardial regeneration via a vascular route (intravenous or intracoronary), has not been as successful as direct injection into the infarcted myocardium. Furthermore, comparative high volumes are needed [98, 100].

Cardiovascular imaging with IONP and MRI

IONPs have also demonstrated capabilities in imaging cardiovascular pathologies including atherosclerosis, thrombosis and myocardial infarcts [56, 101].

Atherosclerosis is considered both a progressive disease and chronic inflammation. The endothelial cells of the vascular wall express receptors from their cell membrane to attract monocytes. Monocytes establish themselves in the subendothelial space and then differentiate into macrophages. The macrophages then take up oxidised low density lipoproteins. Therefore, this lipoprotein can be conjugated to IONPs and used to identify active regions of atherosclerosis with MRI. This approach has been often used in studies involving animal models [102]. Furthermore, apoptosis leads to plaque instability and is known to occur before plaque rupture [102]. Identifying apoptic plaques in vivo with MRI is seen as advantageous in improving patient outcomes, as inflammatory activity of atherosclerotic plaques may be associated with an increased risk of rupture [98].

Aortic valve disease is also an inflammatory response with macrophage involvement [75]. Ruehm et al. [103] successfully used commercially available IONPs to demonstrate (with MR imaging) in hyperlipidemic rabbits that atherosclerotic plaques containing macrophages take up these nanoparticles. This was confirmed with histopathogical evaluation techniques on samples removed from the aorta. The MRI appearance demonstrated an aorta with irregular pattern where the signal dropout occurred at plaque sites containing macrophages endocytosing the IONPs.

IONPs can also be used to target fibrin containing thrombi. This approach may be considered a sensitive method of identifying patients at risk of more serious cardiovascular consequences [104]. A study by Winter et al. [105] demonstrated that by using antifibrin antibodies together with nanoparticles, clots in canine plasma in vitro could be identified with MRI (at 4.7 Tesla).

Blood pool contrast agent

There has been a high level of success in using IONPs as blood pool agents; in both animal models and in human subjects.

One study [106] investigated size and dose of IONPs as a blood pool agent for MR angiography in New Zealand White rabbits. The rabbits were divided into a control group (administered with gadopentate dimeglumine), and three other groups (each receiving a different concentration of IONPs). MR angiography was performed at varying time points; ranging from first pass to 24 hours post-intravascular administration. Assessment included signal enhancement of the abdominal aorta, renal arteries and the iliac arteries. Results demonstrated that the highest level of signal enhancement was identified during the first pass imaging strategy. It was also noted that enhancement in the abdominal aorta was greatest with the smallest nano particle size of 21nm. In addition to this, a concentration of 40micromole of iron per kilogram was recognised as the dose providing signal enhancement comparable to gadopentate dimeglumine. The size and dose of IONPs provided sufficient signal changes to allow imaging from first pass time-point up to 25 minutes post-intravascular administration. In this 25-minute window, multiple imaging acquisitions and measurements are possible.

Another study [107] successfully demonstrated the use of IONPs as a blood pool agent in rats with induced myocardial injuries. Here, a commercially available blood pool agent, Clariscan, was used with an MRA technique. The image findings were also compared with post-mortem histochemical stains of the infracted rat myocardium. It was found that the Clariscan-enhanced MR angiogram images over-estimated the infracted myocardial regions when compared with histology inspection. The use of Clariscan identified the presence and dimensions of both transmural and non transmural microvasculature insults. The peri-infracted zone was seen to be over-estimated to a greater extent in rats with non-transmural ischaemic injuries. Overall, there was an over-estimation of the size of the true infarct but an under-estimation of the region at risk.

Another commercially available IONP preparation (Resovist) has been used to assess the abdominal aorta and the inferior vena cava in a human pilot study [108]. With a 3D MRA (T1-weighted) acquisition technique at 1.5 Tesla, first pass imaging was achievable with results being comparable to those obtained by MRA with conventional Gadolinium based contrast agents.

Ferumoxytol has also been used for first pass enhanced 3D MRA of blood vessels of varying dimensions in 12 human subjects. The following vessels were investigated: the carotid arteries, thoracic aorta, abdominal aorta and peripheral arteries [109]. With delayed MRA imaging it was noted that both arterial and venous structures displayed enhancement characteristics.


A critical and relatively new-found application of IONPs is to include them in preparations to produce what is now termed �nanosensors�. Nanosensors, in their broadest definition, are IONP preparations that have been designed to detect the presence of a specific biological interaction.

Protease specific nanosensors have been developed for in vivo detection of enzyme activity with MRI [110]. In particular, T2* MRI together with nanosensors of 25 nm (hydrodynamic size) have demonstrated the involvement of the metalloproteinase 9 (MMP-9) enzyme in processes including inflammation, atherosclerosis and tumour spread. This opens the potential for early diagnosis of pathologies with altered protease activity and also the monitoring of treatment and therapies that act on protease enzymes.

Protein functionalised nanosensors have also demonstrated capabilities in identifying and measuring the concentration of anti-human serum albumin antibodies [111]. One of the additional benefits of the combined approach of MRI with nanosensors is that the NMR capabilities can allow for detection in natural human substances such as blood and urine.

Nanosensors have also been developed with the capability to detect single human alveolar cancer cells (A549) within 15 mins in blood in vitro [112]. High density folic acid, when conjugated to polyacrylic acid coated IONPs were designed to interact with A549 cancer cells expressing the folate receptor.

Metal doped IONPs

Metal doping of IONPs is designed to provide a comparatively higher level of magnetism at the nano scale and also allow successful magnetic tuning [113]. Such metal doped IONPs can increase MRI signal contrast by as much as 14 times compared to conventional IONPs. The implication of this is that a lower dose or lower concentration can be administered to the patient. Their magneto-thermic effects can also increase by a factor of 4.

Lanthanide metals have also been used to dope IONPs [114]. The advantages that lanthanide metals can offer include optical imaging properties, detection by neutron activation, utilised in neutron capture therapy procedures and also being detectable by time resolved fluorescence.

IONPs doped with manganese have shown to improve the quality of contrast-enhanced MR imaging of the liver [115]. These nano systems have a mean diameter of 80nm. The highest level or change in signal intensity occurs at 5 mins post-intravascular administration. However, unlike early IONPs, the imaging window for IONPs doped with manganese can last for approximately 36 hours.


In the immediate future, the reviewers see that perhaps the most likely application of IONPs will be as an MRI contrast agent. This is because IONPs offer several important advantages over commonly available gadolinium based contrast agents; including, but not limited to, lower toxicity. They can be used to diagnose a variety of pathologies and biological activities such as inflammation, infarction and tissue repair.

In addition to improved diagnostic outcomes, there can also be tailored therapeutic functions, in particular with tumour treatment. IONPs can be binded with antibodies, drugs, enzymes and proteins. They can be directed to specific organs or tissues and can also be guided with magnetic fields to target tumours and induce hyperthermic effects.

Multifunctional IONPs, together with MRI, offer unique advantages with diagnostic and therapeutic capabilities. In particular, molecular imaging with IONPs and MRI has the potential to heavily impact upon early detection of a variety of diseases and pathologic processes. It can also be used to devise treatment approaches dedicated to individual patient circumstances. This should all lead to improved patient outcomes. These are areas of high research activity and there is no doubt that this will lead to a new frontier in magnetic resonance imaging.

Figure 1 Schematic representation of a basic iron oxide nanoparticle, or nanosystem.


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Received 14 October 2009; received in revised form 22 December 2009; accepted 25 January 2010

Correspondence: Division of Medical Radiations, School of Medical Sciences, RMIT University, Bundoora West campus, PO Box 71, Bundoora 3083, Victoria Australia. Tel.: + 61 3 9925 7908; Fax: + 61 3 9925 7466; E-mail: (Giovanni Mandarano).

Please cite as: Mandarano G, Lodhia J, Eu P, Ferris NJ, Davidson R, Cowell SF, Development and use of iron oxide nanoparticles (Part 2): The application of iron oxide contrast agents in MRI, Biomed Imaging Interv J 2010; 6(2):e13

University of Malaya, Kuala Lumpur, Malaysia


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