Cardiovascular magnetic resonance imaging (MRI)

Cardiovascular magnetic resonance imaging (MRI)

8 Cardiovascular magnetic resonance imaging (MRI) R. M. B O T N A R, King’s College London, UK and M. R. M A K O W S K I, King’s College London, UK an...

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8 Cardiovascular magnetic resonance imaging (MRI) R. M. B O T N A R, King’s College London, UK and M. R. M A K O W S K I, King’s College London, UK and Charité University of Medicine, Germany DOI: 10.1533/9780857097477.2.214 Abstract: Cardiovascular disease is the leading cause of morbidity/ mortality in the western world. Atherosclerotic plaques are the main culprit of myocardial infarction and stroke. Magnetic resonance imaging is an emerging clinical tool for the detection and characterization of atherosclerotic plaques. Target-specific molecular MR contrast agents allow the characterization of plaques on a molecular and cellular level. This chapter starts with an overview of the biological changes leading to atherosclerosis. The basic principles of molecular MR are discussed and different MR sequences are introduced. This chapter concludes with a discussion of clinically-feasible aortic, carotid and coronary MR imaging techniques. Key words: cardiovascular magnetic resonance imaging, atherosclerosis, contrast agents.



Cardiovascular diseases (CVD) remain the leading cause of morbidity and mortality in the western world and developing countries. Diseases affecting the cardiovascular system are expected to be the main cause of death globally within the next 25 years. This is due to the rising prevalence of CVD in the eastern parts of Europe, developing countries and an increasing incidence of obesity and diabetes in the western world (Rosamond et al., 2007). Coronary arteries with atherosclerotic plaques are associated with most events resulting in cardiovascular mortality and morbidity. In clinical practice, the early identification and characterization of atherosclerotic lesions that cause myocardial infarctions (MI) and other complications is difficult. Invasive X-ray coronary angiography is considered the ‘goldstandard’ for the diagnosis of significant coronary artery stenoses. Although various non-invasive procedures are available to help identify those with and without significant coronary artery stenosis, up to a third of patients referred for X-ray coronary angiography do not require interventional treatment (Budoff et al., 1996). Severe atherosclerotic vessel wall changes may be missed on X-ray angiography, as positive remodeling may occur 214 © 2014 Woodhead Publishing Limited

Cardiovascular magnetic resonance imaging (MRI)


without significant luminal narrowing (Glagov et al., 1987). Atherosclerotic plaques leading to positive vessel wall remodeling may have a higher risk of plaque rupture with its complications (Rentrop, 2000). The main complication after atherosclerotic plaque rupture is thrombosis of a coronary artery with subsequent vessel occlusion, resulting in an acute myocardial infarction (Rentrop, 2000). It has been suggested that multiple potentially vulnerable lesions can be present simultaneously in the same coronary artery in patients who experience a myocardial infarction (Libby, 2002). Cardiovascular MRI (CMR) is an emerging non-invasive imaging modality and MRI of the arterial vessel wall allows the non-invasive characterization of changes associated with atherosclerosis (Botnar et al., 2000). In addition, molecular MRI may provide useful insights into biological events at a cellular and molecular level in vivo, and has the potential to improve the assessment and characterization of normal and disease states of the arterial vessel wall. The first part of this chapter discusses the technical challenges of and the imaging strategies available for arterial vessel wall imaging of the aorta, the carotid and coronary arteries. Following this, visualization of the arterial vessel wall on a molecular level using targetspecific contrast agents is discussed.


Biology of atherosclerosis

Pathological alterations of the vessel wall in atherosclerotic disease occur mainly in the arterial intima. This process involves the infiltration of modified lipids such as low density lipoproteins, macrophages, smooth muscle cells and the expression of different cytokines and extracellular matrix proteins (Libby, 2002). Inflammation plays a central role in the initiation and progression of atherosclerosis. Low-density lipoproteins infiltrating and being retained in the intima are considered the precursors of atherosclerosis (Libby, 2002). Once LDL is oxidized, it activates endothelial cells that subsequently express adhesion molecules such as E-selectin and vascular cell adhesion molecule-1. The role of adhesion molecules is to attract inflammatory cells, leading to the migration of monocytes and T-lymphocytes from the blood stream into the intima. Within the intima, monocytes differentiate into macrophages and phagocyte modified LDL particles (Fig. 8.1). This process subsequently leads to a transformation of macrophages into foam cells and the accumulation of cholesterol. At this stage of development, these lesions can be classified as so called ‘fatty-streaks’ (Libby, 2002). Subsequently, SMCs originating from the media start migrating into the intima and begin to proliferate. These proliferating SMCs constitute a substantial part of the intima and are primarily responsible for the expression of extracellular matrix proteins such as elastin and collagen (Krettek et al., 2003). In addition, macrophages


Biology of Atherosclerosis Thrombus


Fibrous cap Monocyte




Endothelial cell

Necrotic core

, tins lec 1, e S AM IC AM1 VC



Foam Cell Cytokines, LDLMacrophage Growthfactors





EEM Smooth muscle cell Endothelial Endothelial dysfunction activation Selectins Integrins



ECM formation



Collagen, Elastin

8.1 Different stages in the development of atherosclerosis.

Lipid core, Thrombosis Fibrous cap (Fibrin) formation Fibrin, Lipid core, Fibrous cap Platelets

Cardiovascular magnetic resonance imaging (MRI)


express pro-inflammatory cytokines, matrix metalloproteinases (MMPs) and tissue factors. The increasing accumulation of lipids within the intima leads to the formation of lipid-rich necrotic cores. As a compensatory mechanism to accommodate the expansion of the intima, arteries can enlarge, which is typically referred to as outward- or positive-remodeling (Glagov et al., 1987). If inflammatory processes persist and the infiltration of modified lipids continues, the lipid core continues to enlarge, while proteinases may degrade extracellular matrix proteins. These processes can render the fibrous cap thin and make it susceptible to rupture (Krettek et al., 2003). If rupture occurs, coronary blood can get into direct contact with exposed tissue factors, thereby triggering the coagulation cascade. Thrombin converts fibrinogen to fibrin and plays a role in activating platelets, subsequently initiating thrombus formation. If newly-formed thrombi occlude a coronary artery, an acute myocardial infarction can be the consequence. In some cases, thrombus can be resolved by endogenous thrombolysis (Libby, 2002).


Principles of cardiovascular magnetic resonance (MR) imaging

8.3.1 Imaging contrasts in MRI Magnetic resonance imaging is a non-ionizing tomographic imaging modality, which can be used to generate anatomical and functional images with excellent soft-tissue contrast of the human body. The nuclear magnetic resonance phenomenon was discovered by Bloch, Purcell and co-workers (Bloch, 1946; Purcell et al., 1946) in 1946. 25 years later, its first potential application was reported (Lauterbur, 1973; Mansfield and Grannell, 1973). This chapter is an introduction to the basic principles of nuclear magnetic resonance. A focus is on topics relevant to the following chapters. The generation of a measurable signal in MRI is based on the absorption and subsequent emission of energy in the radio frequency range. Due to its high natural abundance in the human body, hydrogen 1H is used for imaging in a medical context. The hydrogen atom consists of a single electron and proton. Atomic nuclei having odd atomic numbers and/or atomic weights, (e.g. hydrogen atoms) possess an angular moment, known as spin. Nuclei with both even mass and charge numbers do not exhibit spin angular momentum. This property can be displayed as a spinning motion of the nucleus about its own axis (Fig. 8.2). Associated with the spin angular momentum is the magnetic property of the atom. This microscopic magnetic field can be represented by a vector, called magnetic moment. Due to thermal random motion, the direction of the magnetic field of each 1H is random in the absence of an


Biomedical Imaging z B0 w0 m


8.2 If an externally magnetic field B0 is applied, the orientations of the protons (magnetic moments) align at a specific angle along with or opposed to the field. The high- and low-energy states are referred to as ‘spin down’ and ‘spin up’. Adapted from Gadian et al., 1995.

external magnetic field.Therefore, no measurable spontaneous magnetisation exists around a macroscopic object. If an external static magnetic field B0 is applied along the z-axis, the nuclei acquire energy E due to the interaction between the field B0 and the nuclear magnetic moment. The orientations of the magnetic moments align at a specific angle along with or opposed to the applied external field, known as Zeemann splitting. The two opposing orientations display slightly different energies, with the difference being proportional to the magnitude of the applied field. By applying an additional magnetic field B1, transitions between these states can be induced. This external field is applied perpendicular to the direction of B0. As an alignment along rather than opposed to the B0 field corresponds to a lower energy state, more nuclei will align along the B0 field, which will yield a net magnetisation along the z-axis, the longitudinal magnetisation. As the magnetic moments will align at a specific angle (±54°) along B0, the resulting torque leads to a precessional movement of the magnetic moments around the B0 field. The Larmor constant characterises this angular frequency(ω0). To sum up the macroscopic magnetisation of a spin system, a vector M is introduced. However, magnetisation along the z-axis cannot be detected; therefore it is necessary to tilt the magnetisation from the z-axis towards the xy-plane. This can be achieved by a radiofrequency field B1, applied in the xy-plane.

Cardiovascular magnetic resonance imaging (MRI)


For a better understanding and visualisation of the effects of B1, the concept of a rotating frame of reference, spinning about the z-axis at frequency ω0 is introduced. In this reference system the nuclear magnetisation can be described as stationary. RF-pulses of B1 will shift the net magnetisation away from the z-axis into the xy-plane of the rotating frame. Hence, changes in the net nuclear magnetisation can be caused by applying an oscillating B1 field with an identical Larmor frequency to the nuclei.  The motion of M in the presence of a strong external magnetic field B0 and the magnetic component B1 can be described by a set of differential equations (Bloch, 1946).

ω 0 = γ B0


This equation shows that magnetisation will precess around a magnetic field at the rate ω. This rate is proportional to the strength of the field. To describe the motion of M0 more precisely, the different ways in which nuclear relaxation occurs must be described, namely spin–lattice (longitudinal) and spin–spin (transverse) relaxation. In the process of longitudinal relaxation M0 reverts to its original state along the z-axis; in the process of transverse relaxation the M0 component in the xy-plane decays. The result of spin–lattice (longitudinal) is to bring the M0 back to its original equilibrium state along the z-axis, if it is perturbed (Gadian, 1995) and is described by a time constant T1. This is termed spin–lattice relaxation, as exchange of energy between the nuclear spins and their molecular framework takes place in the process. To explain the phenomena that are responsible for relaxation, it has to be considered that nuclei are within a lattice structure and are in constant motion, creating a complex magnetic field. Nuclei in a higher state of energy can transfer their energy to nuclei in a lower energy state as molecular motion fluctuates at the resonance frequency. This phenomenon is called dipole–dipole relaxation (Gadian, 1995). This relaxation mechanism is most efficient when the frequency of the molecular motion is equal to the resonance frequency

τ c = 1/ω 0 (τ c is the correlation time).


Transverse relaxation results from spin–spin interactions and is described by the time constant T2 (transverse relaxation time) and leads to the return of the transverse magnetisation Mxy to its equilibrium. Interactions between adjacent nuclear spins with identical precessional frequencies but different magnetic quantum states, allow for the exchange of quantum states and result in transverse relaxation. An excited nucleus at a higher energy state can excite a nucleus at a lower energy level, thereby relaxing to a lower energy state.


Biomedical Imaging Mxy (t ) = Mxy e − t /T2


This equation reflects the exponential decay of the transverse magnetisation to zero with the tissue-specific relaxation time T2. This process can also be described as decoherence of the transverse nuclear spin magnetisation. If an excitation RF pulse is applied, magnetisation is shifted into the xy-plane and the different nuclei precess at slightly different frequencies. The greater the frequency spread between the different nuclei, the faster Mxy decays. 1 1 = + (inhomogeneity effects) T 2* T 2


If both T2 relaxation and the effect of field inhomogeneities are considered, the FID generated by an RF pulse decays at a faster rate compared to what is expected by the T2 relaxation. This decay is characterized by a time constant called the T2* relaxation time, which describes the tissue T2 value and the contribution from field inhomogeneities. Unlike the T1 and T2 relaxation times, the T2* time is not an intrinsic property of a tissue. In most cases, the additional inhomogeneity, and thus the value of T2*, are dependent on the distribution and magnitude of field inhomogeneities and on the size and shape of the voxel. As long as the field inhomogeneities do not change over time, the dephasing, and the associated portion of the MR signal decay, is reversible. In summary, transverse and longitudinal relaxation processes are caused by fluctuating magnetic fields. Slow molecular motions contribute solely to spin–spin relaxation, components of molecular motion at the resonance frequency contribute to both transverse and longitudinal relaxation (Gadian, 1995).

8.3.2 Properties of MR contrast agents The aim of the first part of this chapter is to provide an overview of the basic principles of molecular cardiovascular MRI imaging. Different groups of contrast agents are introduced and the current state of cardiovascular molecular imaging with targeted contrast agents is reviewed (Fig. 8.3). This chapter will start with a focus on physical principles of molecular MRI, while the later part concentrates on basic properties of targeted contrast agents, including target identification, targeting approaches and signal amplification strategies. At the end of this chapter, current and potential applications of molecular MRI in atherosclerosis are described and discussed. Cardiovascular molecular MRI with targeted contrast agents refers to the non-invasive imaging of biological processes at a molecular and cellular level to improve the in vivo detection and characterization of

Cardiovascular magnetic resonance imaging (MRI)


Small molecular weight CAs

Nanoparticles (10nm–1mm)

A) Rapid clearance (renal) B) Rapid distribution in extracellular space C) Moderate relaxivity (high abundance targets)

A) Slow clearance (liver, cellular uptake) B) Slow distribution in extracellular space C) High relaxivity (high abundance targets)

Carrier Cells Liposomes Microbubbles Perfluorocarbon emulsions Crosslinked iron oxide (CLIO) Target Adhesion molecules Intergins Receptors Fibrin

Signal element Gd3+ chelates or iron oxide Ligand Monoclonal antibodies or fragments Small peptides Free Water

Gadolinium T1 lowering effect→hotspot Sensitivity: mM-mM range Iron oxide T2* effect→signal void (dark) Sensitivity: mM-nM range

8.3 Schematic of molecular MRI contrast agents. The main components of molecular contrast agents consist of a specific ligand that binds to a target and a signal element, which in the case of MR, is made of a Gd3+ chelate or an iron oxide. These two mainly-used basic components can be directly linked to each other or may be attached to or incorporated within a larger nanoparticle (= carrier). Adapted from Jansen et al., 2009.

diseases. Compared to other imaging modalities, MRI can provide high spatial resolution with unique soft-tissue contrast, and the capability to simultaneously image cardiovascular anatomy, function and changes at a molecular level (Johnson et al., 1993). In contrast to other imaging modalities, e.g. single photon emission computed tomography (SPECT) and positron emission tomography (PET), MRI has a lower sensitivity for contrast-agent detection, which limits its use for imaging of low-abundant molecular markers. There are different categories of molecular MR contrast agents, e.g. small molecular weight contrast agents and nanoparticles. MR contrast agents can be used to shorten local T1- and T2- relaxation times. MRI contrast agents such as Gd-DTPA (gadolinium diethylenetriamine-penta-acetic-acid) cause


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an increase in the longitudinal relaxation rate (R1), due to their paramagnetic properties. The increase in R1 is found to be directly proportional to the contrast-agent concentration. This concentration can be calculated by measuring the intrinsic relaxation rate R1 before and after contrast agent injection. Contrast agents, however, also influence and shorten the T2 relaxation time (Caravan et al., 1999). Due to the low abundance of certain molecular imaging targets, longitudinal and transverse relaxivity, r1 and r2, are essential for the design of target-specific contrast agents and will be discussed later. Positive-contrast-MR-probes are mainly gadolinium-based and have a slightly stronger T1 lowering effect compared to T2 (Weinmann et al., 1984). These contrast agents therefore lead to a positive contrast effect, which can be detected as an increase in signal intensity or brightness. Typical r1 and r2 values of currently approved Gd-based contrast agents are in the range of r1 approximately 4 (mM x s)−1 and r2 approximately 5.5 (mM x s)−1. Dependent on the design of the Gd-based contrast agent, the longitudinal relaxivity r1 may decrease with increasing field strength. Negative-contrast-MR-probes are based mainly on iron oxide particles, have a stronger effect on decreasing T2 and cause a negative contrast effect, which is detectable as a signal void (Weissleder et al., 1990). Apart from the effect these particles have on decreasing T2, iron particles also decrease T2* due their effect on the local magnetic field B0. The local accumulation of these particles therefore causes local field inhomogeneities. This additional effect leads to a strong local signal decay. Iron-based contrast agents are therefore often imaged using T2*-weighted imaging sequences. The relaxivities of iron-based contrast agents are significantly higher compared with Gd-based agents, allowing detection with higher sensitivity (Farrar et al., 2008). Signal intensity in MRI primarily depends on local values of longitudinal and transverse relaxation rates of water protons. Usually, a threshold for a local concentration of a contrast agent has to be reached to alter the relaxation rate of water protons sufficiently for detectable signal effects. Most MR contrast agents are based on either gadolinium complexes (Laniado et al., 1984; Caravan et al., 1999) or iron oxide particles (Weissleder et al., 1990). Gadolinium (III) is ideally suited for the development of MRI contrast agents as it has seven unpaired electrons and the symmetry of its electronic states produces an electron spin relaxation time slow enough to interact significantly with neighbouring water protons (Caravan et al., 1999). The relaxivity is determined by a number of properties, including hydration number, distance between ion and solvent proton, solvent exchange rate, electronic relaxation time, and rotational correlation time. The ability of bound protons to rapidly exchange with free water results in the distribution of relaxation effects throughout surrounding water. The design of molecular MRI probes can involve the attachment of Gd complexes to small ligands (e.g. small peptides) that bind

Cardiovascular magnetic resonance imaging (MRI)


to specific targets (e.g. receptors). This process results in the lengthening of the rotational correlation time, which amplifies the detection of contrast agents. To amplify the MRI signal for the detection of very low abundant targets, multimer approaches can be used to increase the sensitivity of gadolinium-based MRI contrast agents (Caravan, 2006). The receptorinduced magnetisation enhancement also leads to an increase in the targetto-background signal ratio (Nivorozhkin et al., 2001). An example of this type of contrast agent is gadofosveset trisodium, an intravascular contrast agent. It reversibly binds to albumin in plasma. If bound, its relaxivity increases significantly (Caravan et al., 1999). A different approach to enhance the T1 effect of molecular probes is by increasing the number of gadolinium atoms per targeting complex. This effect can be exploited for both gadolinium chelate–antibody probes and nanoparticle probes. The magnitude of the effect, and thereby the increased signal level, is substantially higher for nanoparticles. This principle appears to apply only to a small fraction of gadolinium atoms in these complexes, which are located at the surface of the particle and therefore able to interact with surrounding water molecules. A drawback of this type of contrast agent is the large size (approximately 10 to 500 nm), reducing the ability to penetrate the endothelium and increasing the blood clearance time. For the generation of negative contrast on T2 and T2*-weighted images, the synthesis and application of nano-sized iron oxide particles has been investigated (Weissleder et al., 1990; Ferrucci and Stark, 1990). Depending on their size, iron oxide particles have different effects on 1/T1 and 1/T2. Superparamagnetic iron oxide particles have a much stronger effect on 1/T2 and are best visualized by T2- or T2*-weighted sequences (Ferrucci and Stark, 1990). In addition, superparamagnetic iron oxide particles lead to a marked disturbance of the surrounding magnetic field, resulting in a frequency shift and static susceptibility gradient leading to additional spin dephasing (stronger signal void). Ultra-small superparamagnetic particles of iron oxide have a stronger effect on 1/T1 and 1/T2, compared with gadoliniumbased contrast agents, and therefore can be also detected by T1-weighted sequences (Small et al., 1993). An advantage of iron oxide-based agents is the higher relaxivity per metal atom compared to gadolinium-based agents, resulting in a higher sensitivity for detection. Gadolinium based agents, however, provide positive T1 signal enhancement over a larger contrastagent concentration range and can be more easily discriminated from potential artifacts. Criteria to select molecular targets include the biologic and clinical relevance, and the feasibility of target identification, which is dependent upon factors such as target density and accessibility to probes. Once a molecular target has been selected, appropriate candidate ligands must be identified. Following identification of an appropriate targeting ligand, it has


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to be attached to a suitable signal element. Usually, the signal element is either a metal chelate or a nanometer-sized particle or liposome with associated signal elements (Gd or iron oxide particles). To increase the avidity of molecular probes to specific targets, multiple ligands can be attached. This process can, however, alter the specificity and affinity of contrast agent to the target molecule. Therefore, in vitro binding studies are essential to assess whether specific ligand–target binding has been altered. A specific receptor, molecule or cell type can be targeted by a contrast agent by either passive or active means. Contrast agents can either bind selectively to molecular targets (active targeting) or, due to specific distribution characteristics, favour particular tissues or cell types (e.g. passive targeting).


Preclinical molecular imaging in atherosclerosis

The infiltration of modified lipids into the intima is one of the early processes of atherosclerotic plaque development (Libby, 2002). An example of a lipophilic, gadolinium chelate is gadofluorine. This contrast agent has a longer circulation half-life and is more lipophilic compared to conventional chelates. In a rabbit model of atherosclerosis, high accumulation of gadofluorine was observed in lipid-rich plaques (Sirol et al., 2004). Recent publications have also shown co-localisation of this contrast agent with areas of collagenous (fibrous) material (Ronald et al., 2009) and with neovessel-rich regions, another feature of plaque vulnerability (Sirol et al., 2004). Other early markers of atherosclerosis include endothelial cell adhesion molecules, such as vascular cell adhesion molecule-1, selectins and intercellular adhesion molecule-1. These proteins play a role in endothelial activation and subsequent influx of inflammatory cells into the subendothelial space (Cybulsky and Gimbrone, 1991). In-vivo imaging of cell adhesion molecule-1 has been successfully achieved in pre-clinical models and uptake of a cell adhesion molecule-1 nanoparticle was measured with T2*-weighted CMR (Nahrendorf et al., 2006). The inflammatory response in atherosclerosis involves the influx of inflammatory cells into the vessel wall (Libby, 2002). A high macrophage content is a characteristic feature of vulnerable atherosclerotic plaques (Libby, 2002). Molecular MRI can be used to visualise macrophages. The most evaluated approaches in experimental models and patients are based on iron oxide particles (Kooi et al., 2003). These nanoparticles are accumulated in macrophages in atherosclerotic plaque. The earliest successful demonstration of non-invasive macrophage imaging was by Ruehm et al. in a rabbit model of atherosclerosis (Ruehm et al., 2001). Iron oxide particles can be detected in vivo by MRI as they shorten the

Cardiovascular magnetic resonance imaging (MRI)


local T2 and T2*-relaxation time. The main limitations of negative-contrast approaches are that it can be difficult to detect and distinguish from surrounding tissues (Farrar et al., 2008). Several positive-contrast techniques have demonstrated unique potential to improve the sensitivity for in vivo detection of iron oxide particles (Liu et al., 2008). Three positive-contrast methods (SGM, GRASP and IRON) were specifically tested in the context of atherosclerosis (Korosoglou et al., 2008; Mani et al., 2006; Liu et al., 2008 and Makowski et al., 2011a). All techniques allowed detection of ironoxide-labeled cells as bright spots (Fig. 8.4). IRON-MRI in conjunction with iron oxide particles was shown to be a promising approach for the noninvasive evaluation of macrophage-rich plaque. During the development of atherosclerosis, extracellular matrix proteins, such as elastin, are mainly expressed by smooth muscle cells (Krettek et al., 2003). A novel elastin-specific MRI contrast agent was recently investigated in the context of atherosclerosis (Makowski et al., 2011b). A good correlation between in vivo and ex vivo plaque burden measurements could be established. In a swine model, this elastin-specific contrast agent was successfully applied for the measurement of coronary vessel wall remodeling (von Bary et al., 2011). Fibrin is an important contributor to the formation of atherosclerotic plaque. It is present mainly in the later stages of plaque development and can be detected in fibroatheromas and thin cap fibroatheromas (Tavora et al., 2010). It is also one of the key proteins involved in thrombus formation after plaque rupture (Libby, 2002). Disruption of the fibrous cap of a vulnerable plaque leads to the exposure of the lipid core and matrix to circulating blood. This results in platelet adhesion to collagen via the GPVI collagen receptor, followed by platelet activation and conversion of fibrinogen to fibrin. Aggregated platelets combine with fibrin to form a thrombus, which can occlude the vessel. In recent studies, fibrin-specific small peptides have been successfully used for imaging of thrombus in the jugular vein (Flacke et al., 2001), aorta (Botnar et al., 2004a, 2004b), the pulmonary (Spuentrup et al., 2005) and coronary arteries (Botnar et al., 2004a, 2004b). This agent has been already used successfully to image thrombosis in clinical patients, thereby demonstrating the feasibility of clinical translation (Spuentrup et al., 2008).


Clinical imaging of atherosclerosis

Cardiovascular magnetic resonance imaging (CMR) has been applied in different studies to assess atherosclerosis in different vascular beds like the aorta, the carotids and the coronary arteries. It was shown to have potential for atherosclerotic plaque detection and characterization in this vascular bed in vivo.

Biomedical Imaging Iron oxide SGM




8 we HFD



12 we HFD


8.4 Assessment of intraplaque macrophages using susceptibility gradient mapping. Imaging was performed in ApoE−/− mice on a high fat diet. In the control group, no signal void could be observed after iron oxide injection. Between 8 and 12 weeks of high-fat diet, a significant increase in intraplaque iron oxide accumulation could be observed as a signal void on T2* weighted images and as positive signal on susceptibility gradient images. Histology confirmed the imaging results. Positive contrast provided by SGM-MRI allowed for a clear visualization of intraplaque iron oxide depositions, and magnitudes (mT/m) of contrast enhancement in SG parameter maps enabled for the quantification of intraplaque iron oxide particles. Adapted from Makowski et al., 2011a.

Cardiovascular magnetic resonance imaging (MRI)


Native T1-weighted, T2-weighted or proton-density-weighted CMR of the vessel wall imaging enables the assessment of wall thickness, area and plaque volume. Measurements of plaque burden can be of particular importance in detecting the early formation of atherosclerotic plaques (Alizadeh Dehnavi et al., 2007). Early disease can be detected by assessing the arterial wall thickness, comparable to intima-media thickness measurements by ultrasound (Zhang et al., 2003). Plaque components can also be differentiated using CMR. The necrotic core usually consists of a combination of lipids, water and potentially blood degradation products. CMR signal intensities on different sequences represent those variations. In most cases, the necrotic core is hypointense on T2-weighted images and isointense on T1-weighted images (Saam et al., 2005). For the visualisation of the fibrose cap, black-blood sequences (T1-weighted, T2-weighted or proton density-weighted) can be used (Kampschulte et al., 2004). Intraplaque hemorrhage is believed to result from leaky neovessels, erosion or plaque rupture. This has been suggested to play a role in plaque progression, as degradation of red blood cells and their lipid-rich-membrane leads to the accumulation of lipids (Virmani et al., 2005) and may represent a risk factor for plaque destabilisation (Singh et al., 2009). Carotid CMR can detect intraplaque hemorrhage due to the short T1-relaxation-time of methemoglobin, resulting in hyperintense signal on T1-weighted images (Moody et al., 2003; Takaya et al., 2006; Altaf et al., 2008). Using carotid vessel wall CMR, it has also become feasible to follow changes in atherosclerotic plaque composition over time. Carotid vessel wall CMR has also been used in combination with extravascular contrast agents. Using these agents, information about the size of the necrotic core can be gained (Yuan et al., 2002; Wasserman et al., 2002). Within the atherosclerotic plaque, inflammation is usually most prevalent in the fibrous cap and at the edges of the necrotic core (Moreno et al., 1994). Inflammation has been shown to be associated with angiogenesis. A significant increase in microvessel density was observed with the progression of intraplaque inflammation and a correlation with the incidence of plaque rupture could be found (Moreno et al., 2004). Microvessels, particularly those with high permeability, may also allow inflammatory cells to infiltrate the growing atherosclerotic lesions. Angiogenesis can be measured by dynamic-contrast-enhanced-CMR. Kinetic modelling of signal enhancement of the vessel wall allows assessment of tissue microvascularity and permeability, which are considered potential surrogate markers for the level of inflammation (Kerwin et al., 2006). Compared with the other arterial beds (carotid arteries and aorta), native and contrast enhanced MRI of the coronary vessels is more demanding. Coronary arteries are smaller in diameter and are subject to significant cardiac and respiratory motion. During respiration alone, the movement of


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the coronary arteries may exceed a multiple of the coronary artery diameter. Initially, breath-hold techniques were used, but as breath-hold duration is limited to 15–25 s, spatial resolution is relatively low and precludes high-resolution coronary-vessel-wall-imaging by CMR. Furthermore, diaphragmatic drift can also result in motion artifacts (Holland et al., 1998). Navigators were developed to enable imaging acquisition during normal breathing of the patient (Stuber et al., 1999). Typically, only data from endexpiratory gating intervals are used for image reconstruction. Due to the inherent intrinsic movement of the coronary arteries during the cardiac cycle, ECG synchronization of the data acquisition is also needed. Coronary artery motion is minimal during a brief period of diastases in systole and mid diastole. The latter is typically used for coronary angiography and vessel-wall imaging. The diastolic rest period increases with reduced cardiac frequency. Hence, a cine sequence should be acquired prior to data acquisition, in order to assess the patient-specific trigger delay and acquisition window. Considerable efforts have been made in the recent years to implement and demonstrate the utility of coronary vessel wall imaging. This approach has potential to non-invasively quantify vessel wall thickness and areas using black-blood techniques (Spuentrup and Botnar, 2006). Using black-blood coronary CMR, the coronary lumen appears dark, whereas the surrounding vessel wall has increased signal intensity. In order to achieve optimal contrast between the vessel lumen and wall, the coronary blood signal is suppressed using a double inversion prepulse (Edelman et al., 1991). Comparable to coronary MRA techniques, early in vivo coronary vessel wall imaging studies used two-dimensional breath-hold fatsuppressed fast-spin-echo or free-breathing navigator techniques (Fayad et al., 2000). In-plane resolutions up to approximately 0.5 × 0.5 mm2 could be achieved. However, due to limited coverage of the coronary artery tree, more sophisticated approaches to cover larger portions of the coronary tree were developed. By adjusting the imaging plane to the course of the coronary arteries in combination with modified black-blood prepulses, in-plane views of the coronary artery wall of the proximal and mid portions of the RCA and LAD could be obtained with good contrast between coronary blood and the vessel wall (Botnar et al., 2001). A variety of different pulse sequences have been used to acquire the coronary vessel wall, each with its own advantages and disadvantages. Spiral image acquisition offers the advantage of high SNR and short acquisition windows (Botnar et al., 2001). Free-breathing black-blood 3D spiral coronary vessel wall CMR demonstrated increased coronary vessel wall thickness, with preservation of lumen size in patients with subclinical coronary artery disease, consistent with a ‘Glagov’ outward arterial remodeling (Kim et al., 2002). In a cross-sectional study of more than 100 subjects with longstanding Type 1 diabetes without symptoms or history of cardiovascular

Cardiovascular magnetic resonance imaging (MRI)


disease, MR vessel wall imaging revealed a significantly greater coronary plaque burden and higher prevalence of coronary artery stenosis in patients with nephropathy compared with those with normoalbuminuria (Kim et al., 2007). In another study, black-blood coronary wall CMR allowed detection and quantification of significant positive arterial remodeling in asymptomatic men and women with subclinical atherosclerosis (Miao et al., 2009). Delayed-contrast-enhanced CMR (DCE-CMR) represents an alternative imaging approach for coronary vessel wall imaging by allowing assessing contrast uptake in the vessel wall. Clinically used extracellular nonspecific contrast agents rapidly extravasate into the vessel wall and enhance areas with increased distribution volume. In contrast to native vessel wall imaging, which requires high spatial resolution to assess plaque morphology, requirements on spatial resolution are less stringent as only the presence or absence of contrast agent uptake needs to be detected (Schar et al., 2003). Our group recently investigated the changes of atherosclerotic plaque enhancement in patients with acute myocardial infarction one week and three months after successful coronary intervention. Data suggest that coronary vessel wall contrast uptake is significantly increased early after myocardial infarction as compared to the three-month follow-up scan (Ibrahim et al., 2009). In patients with chronic stable coronary artery disease, the correlation between coronary vessel wall enhancement and luminal narrowing is most likely related to an increased distribution volume for gadolinium due to a fibrous-rich lesion.


Conclusion and future trends

MR vessel wall imaging has matured over the last decade. Important improvements of MR sequences and hardware have been made that allow for the characterisation of the carotid and aortic vessel wall. Various studies have shown that carotid plaque characterisation by CMR in patients with similar degrees of stenosis can differentiate patients with high risk for developing clinical complications from those with a low risk. In specialised imaging centres, coronary vessel wall imaging has been shown to allow non-invasive detection of coronary plaque burden in specific coronary artery segments and the detection of positive arterial remodeling, potentially allowing for a improved risk stratification. Recent studies using molecular MRI techniques showed the feasibility of in vivo elastin, fibrin and macrophage detection in patients with atherosclerosis. Ongoing developments of novel targeting compounds may open new possibilities for biological imaging in atherosclerosis. In conclusion, MRI of the arterial vessel wall has shown great potential to improve clinical risk assessment of patients with atherosclerosis. With further technical developments of MR sequences and contrast agents, CMR


Biomedical Imaging

has potential to be increasingly used for the assessment of atherosclerosis in clinical practice.



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Biomedical Imaging

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Biomedical Imaging

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