A. Introduction

Cardiac magnetic resonance imaging (CMR) has become a reference standard modality for imaging in cardiology. Emerging CMR techniques enable a more comprehensive examination of the heart, making CMR an excellent tool for translational cardiovascular research. Specifically, emerging CMR methods have been developed to measure changes in ventricular mechanics, extent of myocardial edema, fibrosis, and myocardial perfusion as a function of both disease and infarct healing. New CMR techniques also allow cell tracking, molecular imaging of biomarkers of disease, and measurement of changes in cardiomyocyte calcium flux. MRI can also quantify blood flow velocity and wall shear stress in large blood vessels. Almost all of these techniques can be applied in both pre-clinical and clinical settings, enabling both the techniques themselves and the knowledge gained using such techniques in pre-clinical research to be translated from the laboratory bench to the patient bedside. Diseases of the heart, including myocardial infarction (MI) and heart failure, continue to be a leading cause of morbidity and mortality in the western world. In the United States alone, over 5 million people receive medical care for heart failure each year, at a cost of tens of billions of dollars (1). Research into the mechanisms of and potential therapies for cardiovascular disease has increasingly relied upon the use of animal models, with the ultimate goal of translating discoveries from the lab bench to the clinic. This aim of this section chapter is to highlight and to give to the cardiovascular scientist a clear understanding of the power of MRI in translational research.

B. CMR Measures of Structure and Function for Intermediate End-Points

Cardiac morphology, function and CMR

CMR is not constrained by anatomical imaging windows and so images can be obtained in any desired plane or orientation. CMR provides spatial and temporal resolution of the myocardium and therefore allows precise and reproducible measurements of cardiac dimensions. Following initial ‘scout’ images, multiple cine image slices in short axis along the length of the ventricles from base (level of mitral and tricuspid valves) to apex are obtained. By drawing ‘regions of interest’ at end-diastole and end-systole around the endocardium and epicardium accurate measurement can be made of end-diastolic (EDV) (Fig. 1) and end-systolic volumes (ESV), allowing calculation of stroke volume and ejection fraction (EF) as well as myocardial volume (which can be converted to mass by multiplying by 1.05) (2). Therefore for any study looking to accurately quantify changes in-vivo of adverse or reverse ventricular remodelling. CMR is currently the reference standard.

Figure 1. A short axis stack encompassing the whole LV in diastole with the endocardium traced in order to calculate the end-diastolic volume. By tracing the endo- and epicardium in systole too the ejection fraction and LV mass can be calculated.

CMR derived morphological indices may be important predictors of future events and thus may serve as useful surrogate trial endpoints (3). In patients with ischemic heart disease CMR derived indexed LV EDV and ESV are associated with all cause mortality with a hazard ratio (HR) of 1.13 (confidence intervals 1.05-1.23; p=0.002 and 1.04-1.22; p=0.004 respectively) for each 10ml/m2 increase and the same study demonstrated that LV EF is linked to all cause mortality with a HR of 1.36 (CI 1.08-1.72; p=0.01) for each 10% decrease.

Quantification of regional function by CMR (Deformation imaging)

The regional assessment of myocardial function enables further stratification of disease rather than expressing ventricular function only as a global ejection fraction. With the advancement in regionally directed therapies (such as injected cell therapies) there is a developing need for accurate, reproducible techniques in order to study the changes in function in the treated area (4) and their impact on global function and clinical outcomes. Deformation imaging by CMR allows qualitative and quantitative measures of regional myocardial function. This is not yet established in routine clinical practice yet it is increasingly being used in research studies in order to characterize observed phenomena (5,6) or to act as surrogate endpoints in clinical studies (7). Time to peak contraction (assessment of dyssynchrony), regional ejection fraction (an expression of the regional contribution to the global function) and myocardial strain may all be quantified.

Strain is a measure of deformation and is defined by the change in length of an object relative to its original length and in the heart this yields radial, circumferential and longitudinal values. Depending on the technique used the measured strain may be based on a specific point in the myocardium (Lagrangian) or on a fixed point in space (Eulerian). When comparing measurements made with different techniques it is important to bear this in mind as values will be different. Strain measurements made between different modalities may also not be comparable, such as those by echocardiographic speckle tracking and CMR myocardial tagging (8). A number of papers have established normal values in reference populations however studied numbers are still relatively small (9,10,11) and expected values in defined disease groups are being formulated (6,12). A number of CMR techniques have been developed and offer varying attributes.

Myocardial tissue tagging

This is the most widely used CMR technique and consists of a preparation phase and then an imaging phase. By nulling the magnetisation in two oblique planes a visible grid (Tags) is produced through the myocardium (Fig. 2) which can then be seen to deform during systole with cine imaging. Progressive technical advancements have improved signal and contrast to noise ratios, increased the duration of the cardiac cycle that can be imaged and reduced the time required for post-processing (notably HARP [Harmonic Phase] analysis).

Figure 2. Diastolic (left) and systolic (right) images of myocardial tagging demonstrating normal myocardial deformation in a volunteer.

Phase velocity encoding

Phase velocity encoding or mapping is most commonly used for assessing flow in large blood vessels but can be used to assess myocardial velocity. A bipolar gradient is used to encode velocity directly into the phase of the signal. Although prone to motion artifacts and phase distortion, phase velocity encoding has been used in a number of clinical studies to measure the relative velocities of the myocardium and hence detect regional wall motion abnormalities. Peak velocity, time to peak velocity and strain rate can all be measured or calculated from the sequence.

Displacement encoding with stimulated echos (DENSE)

DENSE uses a stimulated echo sequence in order to encode in-plane (13) or through-plane (14) displacement directly into the phase of an image. Displacement of each pixel can be calculated either at the extremes of movement or as cine-DENSE where the pixels can be tracked in order to generate strain-time curves. This technique offers rapid processing times and high precision but has not been widely adopted as yet due to a relatively low signal to noise ratio and similar to myocardial tagging techniques the tagged areas fade through the cardiac cycle.

Strain encoded imaging (SENC)

SENC is a further technique which tags the myocardium, but the tags however, are applied in parallel to the imaging plane rather than perpendicular to it. Longitudinal strain can be measured from the short axis views and circumferential strain from the long axis views. Radial strain cannot be measured, although SENC can be combined with other 2D tagging techniques such as 2D DENSE (15) in order to provide a 3D model. Strain is directly related to pixel intensity and therefore post-processing times are short.

C. CMR Tissue Characterization for Intermediate End-Points

By taking advantage of the different behaviours of tissues within a magnetic field, MRI offers the ability to characterize properties of that tissue. By adding a contrast agent such as gadolinium which accumulates in the extra-cellular space (and shortens T1 relaxation) and by imaging at different time points post-administration, further information can be gleaned about myocardial perfusion, microvascular patency, extent of infarcted tissue and potential viability.

Myocardial infarction quantification

The ability to accurately image the size of the permanent myocardial injury that results from a myocardial infarction (either experimentally induced or otherwise) plays a crucial role as a surrogate endpoint in translational science. Gadolinium based contrasts accumulate in the interstitial space and as a myocardial infarction has a large area of interstitial expansion, it is possible to accurately delineate infarct size with excellent correlation using pathological methods (16,17) and clinical techniques (18). The size of CMR determined myocardial infarction is correlated to both functional (19) and clinical outcomes (20).

The peri-infarct zone

The ‘peri-infarct grey zone’ is an area of enhancement which is intermediate between that of infarction and the black colour of ‘nulled’ normal myocardium. It may be partly created by partial voluming due to limitations of the resolution of the CMR technique but it has also been demonstrated that it is the area of the infarction bordering completely non-viable infarcted tissue and completely healthy tissue (21). It is thought that this area provides an ideal substrate to form the re-entry circuits needed for malignant ventricular arrhythmias and that its size is correlated with outcome (3), although whether it provides additional prognostic information to total infarct size is debated (22). This area is however an area of significant remodelling in the post-infarction period and as such its characterization and the ability to chart its modification may be useful surrogate endpoints for interventions aiming to reduce adverse remodelling post infarction (23).

Microvascular obstruction and myocardial haemorrhage

Imaging in the ‘early’ period following contrast administration enables visualisation of any microvascular obstruction (MVO) present from the acute myocardial infarction and is the optimal method of detecting intraventricular thrombus.

MVO is the result of injury sustained to the microvasculature during the period of ischemia and reperfusion or from distal coronary microembolism resulting from atherosclerotic debris. It may be seen as a dark area on images shortly after gadolinium contrast injection (‘early’) or as a dark core within a hyperenhanced area if imaged 10-15 minutes after gadolinium administration (‘late’). MVO may occur despite adequate epicardial coronary artery revascularisation and it is thought to result in the angiographic phenomenon of ‘slow’ or ‘no reflow’ (24). MVO is caused by a number of factors including distal embolization and reperfusion injury with microvascular inflammation and disruption and the activation of vasoactive mediators. The area of MVO is also comprised of myocardial haemorrhage from red blood cell extravasation into the extra-cellular space (25). The presence of MVO and myocardial haemorrhage has been shown to correlate to larger infarct size, increased adverse remodelling and reduced LV function over time following acute myocardial infarction (26,27) (Fig. 3). Optimal reperfusion should aim to restore micro- as well as macrovascular flow. Measurement with MRI allows accurate assessment of strategies aimed at reducing MVO (28). The specific area of myocardial haemorrhage can be imaged using techniques which take advantage of the magnetic properties of degradation products of blood in this region, generally T2 or T2* techniques (29,30) or more recently with T1 weighted techniques (31).

Figure 3. Late gadolinium enhanced imaging demonstrating a recent transmural anteroseptal myocardial infarction with a dark core of microvascular obstruction (arrow).

Area at Risk (AAR)

The area at risk is defined as the area of potential maximal injury that may result from the occlusion of a coronary artery or arteries in the absence of reperfusion (32). This is based on the wave front theory of propagation of infarction which describes the progression over time of the area of infarction from the endocardium towards the epicardium (33). Myocardial salvage is calculated by expressing the final myocardial infarction size as a function of the area at risk (AAR). In the majority of animal experiments the measurement of myocardial salvage is fairly straightforward as the perfused heart can be stained or marked in a number of ways and sectioned to allow accurate quantification of viable and infarcted areas. Besides post-mortem examination techniques it is not possible to measure the AAR in humans in such a direct way and so a variety of techniques are used as surrogates which may limit the translational accuracy. Historically the measurement of AAR has been dominated by nuclear techniques (mainly SPECT (34)) but MRI is attractive as it may provide a longer time window after infarction to perform the scan, avoids radiation exposure and the need to have constant access to radioactive isotopes.

Measuring area at risk (AAR) with the endocardial surface area (ESA) method

The ESA method was described by Ortiz-Perez et al. in 2007 (35) and is relatively easy to perform without the need for extra images beyond those of the infarction and follows on from the ‘wavefront’ theory of infarct progression described by Reimer and Jennings in 1977 (33). The acute infarct is imaged as described above. The lateral extents of the infarct and AAR are set early on in the ischemic period and are delineated by the lateral extent of the gadolinium enhancement. The length of the infarct on the endocardial border is measured, multiplied by the slice thickness to give surface area and this is proportional to the AAR. This is expressed as a percentage of the total endocardial surface area of normal myocardium (calculated in the same way) to give a percentage of LV myocardium at risk. This was originally validated against the two most common angiographic jeopardy scores and performed well (r=0.9 and 0.87, p<0.001) (35). However, a subsequent study did not find such a good correlation particularly in infarcts treated rapidly where the lateral extents of the infarct may not match the potential AAR (36). This technique remains useful but is yet to be proven or adopted on a wide scale.

Measuring area at risk with T2 weighted sequences

Myocardial edema has been recognised to develop rapidly following ischemic injury in as little as 30 minutes (37) and represents an increased water content both intra- and extracellularly. By using T2 weighted short inversion-time, inversion recovery (STIR) imaging CMR sequences (38) it is possible to image the areas of myocardium which have increased water content. In initial animal studies the extent of myocardial edema by CMR corresponded well to quantification of the AAR by fluorescein dye (39) or injection of microspheres (40). Validation in humans against the current gold standard, SPECT imaging, was promising with measurement of the AAR accurate up to 1 week following infarction (41) and the technique has been progressively improved with the development of the cardiac-unified T2 edema (ACUTE) sequence (42) although imaging is still subject to motion artifact and possible misinterpretation. There continues to be disagreement as to whether the area of edema seen with these techniques truly represents the AAR. Recent studies have reported a systematic over estimation of myocardial salvage perhaps due to edematous changes in areas not at risk of infarction (43) and other authors have also urged caution when interpreting T2 edema data (44). The future may lie with the use of sequences which allow quantification or mapping of both the T1 and T2 signal changes induced by edema (45) which is also possible in 3 Tesla scanners (46).

Figure 4. Short axis LV slices following an acute inferior myocardial infarction. The left panel shows a T2 weighted image with bright signal in the inferoseptum and inferior wall (with RV extension) due to myocardial edema, believed to represent the myocardial area at risk. The right panel shows the same segment but in the late phase following gadolinium administration demonstrating a near transmural infarction in the same area although the borders do not quite extend as far as those of the edema.

CMR and edema imaging in other areas

The use of CMR to assess edema or inflammation beyond investigation of the myocardial AAR may prove insightful. The early diagnosis of acute coronary syndrome (47) may facilitate the correct, rapid treatment and may allow recruitment at an early stage to a clinical trial. Edema or inflammation itself may represent disease activity such as in myocarditis (48) or transplant rejection (49).

The diagnosis of cardiac allograft transplant rejection is promising and in the future may provide a means of follow-up in some patients (49,50) permitting non-invasive evaluation of novel therapies. The use of CMR in myocarditis (Fig. 5) is increasing rapidly and is considered essential by some (48). CMR may provide prognostic benefit (51) and may help guide endomyocardial biopsy, although, as yet it cannot be used alone for diagnosis and management (52,53).

Figure 5. A short axis LV slice with a T2 weighted sequence (STIR) demonstrating an area of higher signal intensity in the inferior and inferolateral walls due to myocarditis.

Diffuse fibrosis

Diffuse myocardial fibrosis as distinct from the focal fibrosis of a myocardial infarction is not visualised with the current CMR late gadolinium enhancement technique as the areas of interest are nulled along with ‘normal’ myocardium in order to highlight the focal fibrosis. Diffuse fibrosis of the myocardium is a slow occult process which occurs with normal aging but is accelerated by conditions such as hypertension, diabetes, hypertrophic cardiomyopathy and aortic stenosis (54). Diffuse fibrosis may be reversible and as such it represents a target for novel therapies; therefore a technique that can identify, quantify and repeatedly measure diffuse fibrosis will be crucial for trials in this field (55).

Broadly speaking, two CMR techniques are currently in use to quantitate diffuse fibrosis. Both rely on the fact that fibrosis results in an expansion of the myocardial extra-cellular volume. The first (56) technique evaluates the T1 of the myocardium at baseline and then at different intervals after a bolus injection of gadolinium. The T1 of the fibrotic tissue (with an increased ECV and reduced gadolinium wash out) is shorter than that of ‘normal’ myocardium and by using sequences which map the T1 signal the percentage of ECV versus normal myocardium can be calculated. Histological validation has been conducted in small numbers (57,58). This method is susceptible to inaccuracies at high heart rates and to variations depending on renal function and body composition. A second technique, equilibrium contrast CMR, which requiring greater time and technical expertise, aims to reduce the inaccuracies or variations described above (59). The T1 (of blood and myocardium) is measured before and after gadolinium as previously but in this technique a primed infusion of gadolinium (bolus followed by infusion) is used in order to equilibrate blood and myocardial levels. By taking a blood test to measure the haematocrit the volume of distribution of gadolinium in the blood can be calculated, allowing the calculation of the volume of distribution in the myocardium. Initial validation was performed against human myocardial samples in patients undergoing surgery for aortic stenosis and the group have gone on to investigate the links with clinical outcomes (60).

Clearly the techniques for quantification of diffuse fibrosis are still undergoing rapid advancement but the concept of being able to target diffuse fibrosis as a therapy is a very attractive one.

Perfusion imaging

First pass perfusion imaging

By rapidly imaging the myocardium during an injection of gadolinium contrast an assessment of myocardial perfusion can be made. The most common method is to do this at rest and stress with pharmacological stress provided by an adenosine infusion (61). The use of dobutamine may further allow identification of inducible regional wall motion abnormalities (62) along with a perfusion assessment (63). First pass perfusion imaging with adenosine performs equally well (64), if not better (65,66) compared to nuclear techniques in the diagnosis of ischemic heart disease; and provides important prognostic information (67). With newer techniques whole heart visualisation is possible (68,69), along with regional absolute myocardial blood flow quantification (70). Beyond the diagnosis of relevant epicardial coronary artery stenosis, first pass perfusion CMR also been used to assess microvascular function/dysfunction following myocardial infarction in humans (71) and rodents (72), in hypertrophic cardiomyopathy (73), small vessel vascular disease (74) and poorly understood conditions such as syndrome X (75).

Blood oxygen level dependent (BOLD) perfusion

This technique uses the different properties of deoxyhaemaglobin and oxyhaemaglobin within a magnetic field allowing them to act as intrinsic contrast agents. Deoxyhaemaglobin causes a decrease in T2 or T2* signal and as such can be directly related to oxygenation (76). It has been used successfully in animals to identify experimentally induced ischemia (77) and, with the use of 3T allowing an improved signal to noise ratio, (alongside pharmacological stress) is showing favourable diagnostic results in humans (78,79). It may also be able to give rather different information to first pass perfusion imaging as the level of myocardial oxygenation may not be the same as regional perfusion (80).

Arterial spin labelling perfusion

This is another technique that does not rely on the use of an intravenous contrast. It has significant potential but is still the subject of considerable research in order to overcome the technical challenges of image acquisition (81). Arterial blood is ‘labelled’ by using a slice selective inversion pulse. The non-labelled blood which enters the slice to replace the labelled blood has a shorter T1 and by using this difference between the T1 signals the myocardial perfusion can be calculated. The technique is relatively insensitive at present, covers only small parts of the LV myocardium and is prone to respiratory motion artifacts; however it has demonstrated reasonable accuracy in man (82) and may be an attractive technique in animals to monitor changes in perfusion without needing to administer intravenous contrast (83).

T2* for myocardial iron

Increasing myocardial iron shortens the T2* relaxation time and therefore by calculating the T2* myocardial iron can be quantified (84,85,86). This technique has allowed the diagnosis of siderotic cardiomyopathy and intensification of chelation therapies to prevent heart failure and death (87). Further optimization of the technique may decrease image artifacts and increase its accuracy (88).

D. CMR measures of cell function, cell tracking, and molecular imaging for intermediate End-Points

Calcium channel function/Ca concentration (Manganese enhanced MRI)

In several myocardial diseases such as heart failure (89) or ischemic damage (90), calcium (Ca²⁺) cycling is modified.

Manganese (Mn²⁺) is one of several cations able to selectively block voltage-gated calcium channels. Manganese has a similar structure and kinetic behaviour as calcium (Ca²⁺). In fact, Mn²⁺ also permeates through Ca²⁺ channels, creating a small current (8% of calcium current) (91). Mn²⁺ shortens T1 relaxation time and by this property allows the indirect assessment of calcium flux and cell viability (92).

Manganese enhanced MRI has wide application in the preclinical setting. Validated indications for this method are the determination of myocardial AAR (93) and ischemic areas (94), but also measurement of the effect of ischemia (95) or drugs (96) on the Ca²⁺ current. Safety concerns regarding the physiological and neurological toxic effects of Mn²⁺ (97) have limited its clinical application in man.

Myofiber orientation (Diffusion tensor imaging)

The 3-dimensional organisation of myocardial fibers of the mammalian ventricle is unique; being composed not only of circumferential fibers and longitudinal fibers but also of obliquely running fibers that form a helical spiral from base to apex (98).

Determination of fiber architecture is central to the understanding of certain cardiac diseases and has classically been performed by histological analysis. Diffusion tensor resonance magnetic imaging (DTRMI) has been validated in the late 90’s (99,100). DTMRI takes advantage of the fact that the MRI signal is attenuated by water diffusion in a tissue submitted to a magnetic field gradient. Water diffuses preferentially along the myofibers, rather than across the cell membranes, allowing the determination of the architecture of the myofibers (100) (Fig. 6). Although the technique was initially meant for analysis ex-vivo of myocardium, technical progress has allowed the application of this method to beating hearts in-vivo, making it available for studies in patients (101).

DTMRI’s role in translational research would be to assess the effect of therapies (pharmaceutical, genetic, cellular etc.) on the detailed myocardial structure (98).

Figure 6. Diffusion tensor resonance magnetic imaging has the potential to allow microstructural visualisation and determination of myofiber architecture (102, open access) (courtesy of Dr L-A McGill, The Royal Brompton Hospital CMR Department).

Cell tracking (super paramagnetic iron oxide particles)

Superparamagnetic iron oxide particles (SPIO) are widely used to label cells for MRI detection. The particles contain an iron oxide core surrounded by a polymeric of polysaccharide coat. SPIO produces magnetic in-homogeneities, disturbing the local magnetic field and leading to decreased signal intensity on T2-weighted and T2*-weighted images (103).

The major limitation of using SPIO particles for cell tracking is the absence of discrimination between living labelled stem cells and dead stem cells phagocytised by locally present macrophages. Furthermore, reports indicate than even after 40 days, no signal difference is detectable between labelled living and dead cells (104).

SPIO particles were considered to have no significant interference with the biological properties of the labelled cells (105). However, several recently published papers have demonstrated potential interference with cell migration, colony-formation and cell function (106,107,108).

In humans, intravenous solutions containing SPIO particles have been FDA-approved and several clinical trials in the oncologic and infectious fields have been conducted with these products. Intravenous administration of SPIO-labelled cells was performed recently for the first time in humans and was reported to be safe. The same group also demonstrated that SPIO-labelled cells can be successfully tracked to a target site after systemic intravenous administration (109).

Cells labelled with “classical” MR-contrast such as gadolinium chelates are detectable as a hyperintense signal on T1-weighted images (103). However, the sensitivity of this method is lower than SPIO labelling and, furthermore, contrast intensity is inhomogeneous depending on the intracellular location (110) of the contrast and safety concerns remain in terms of possible release of free radicals (111).

Fluorine (¹⁹F) contained in perfluorocarbon (PFC) is a very attractive emerging MR-contrast. It can be used for spectroscopy and image formation in a similar way as ¹H from H₂O (103). Biological soft tissues do not contain fluorine, preventing any background tissue signal. PFC is known to have a safe toxicity profile as this substance has been commonly used as a blood substitutes for decades. Numerous animal studies have been conducted on stem cell labelling and trafficking with ¹⁹F-MR. Although based on different protocols (cell type, labelling method, delivery route etc), these studies exhibit common results regarding the efficiency of cell labelling, the stability of cell labelling, the specific detection of the cells throughout the body and its quantitative aspect (112). ¹⁹F-MR is currently limited to preclinical research as further improvements are needed particularly in the field of image acquisition (hardware, sequences) and of development of approved ¹⁹F contrast agents before translation to the clinical setting can occur.

Myocardial energetics (MR spectroscopy)

Myocardial metabolism is disturbed in numerous pathologies including ischemia, hypertrophy and failure (113).

In conventional cardiac magnetic resonance, the signal arises from the position of the spins of ¹H nuclei contained in the water and fat molecules related to the magnetic field. This technique provides structural and functional information but is not suitable for metabolic assessment of a particular tissue. CMR spectroscopy allows the study of several other nuclei, including ¹³C, ²³Na and ³¹P being of particular interest for cardiac metabolism (114). In comparison to conventional CMR, MR spectroscopy offers a very low sensitivity related to the low magnetic energy of the nuclear spins and the low concentration of these molecules compared to the abundance of water. To overcome this limitation, the technique of hyperpolarisation has recently been developed. This method uses low temperatures and strong magnetic fields to increase the polarisation of the spins (115).

CMR spectroscopy obviously has a role to play in the understanding of cardiac diseases and in patient management. Although large technological advances have been made, CMR spectroscopy still suffers from the major limitations of poor spatial and temporal resolution, preventing its large-scale application in research or in humans (116). However, hyperpolarized ¹³C-pyruvate has recently been approved for clinical studies on prostate cancer, raising the hope for human studies on myocardial metabolism (117).

E. CMR measures of vascular imaging

Wall stress

The arterial tree is a very complex structure and each segment is submitted to different shear stress conditions depending on its localization. Lowered shear stress has been shown to be strongly associated with formation of unstable plaque (118). MRI studies performed up to 20 years ago tried to derive shear stress from 2D-slices based anatomic reconstruction and flow rate waveforms (119). Recent development of 4D-MRI has been a major step forward in magnetic resonance based vascular imaging since it offers 3D structural visualization combined with time-resolved blood flow velocities in three directions, allowing quantification of blood flow and derived vessel wall parameters (120,121). 4D-MRI is a valuable clinical tool for assessment of vascular geometry and hemodynamics in cases of aortic aneurysms or malformations. Currently, its use in the assessment of vessel wall parameters is limited to research purposes.

Pulse wave velocity

The measurement of pulse wave velocity may give an indication of reduced vascular compliance and of atherosclerosis. Although at present pulse wave velocity measures are rarely used for diagnostic process, they are frequently used for research purposes (122). Furthermore, recent evidence shows that increased aortic stiffness measured by pulse wave velocity is a strong predictor for cardiovascular events and all-cause mortality (123). 4D-MRI-based assessment of pulse wave velocity in the thoracic aorta has been shown to be reproducible in young volunteers and to correlate as expected with age, aortic valve insufficiency and pulse pressure (124). The clinical application of 4D-MRI–based pulse wave velocity measurements is currently for diagnosis and therapy monitoring in particular situations, such as, diabetic children with increased aortic wall stiffness and lipid-lowering therapy (125).

Atherosclerosis (plaque imaging)

Atherosclerosis is a progressive disease where vessel walls are thickened by the development of lipid rich plaques. This process is characterized by cholesterol deposition, inflammation, extracellular matrix formation and thrombosis (126,127). Vulnerable plaques are lesions at risk of destabilization potentially provoking an ischemic event. Features of high rupture risk include a large lipid core, a thin fibrous cap with low smooth muscle cell content, infiltration of inflammatory cells, neovascularization with intraplaque hemorrhage and outward remodeling (128). MRI of atherosclerosis is a valuable tool for diagnosis and follow-up since it offers a non-invasive, ionizing radiation free assessment of plaque volume and composition (126). Three complementary methods are currently available to characterize atherosclerotic plaques. First, the multicontrast method allows visualization of lipid core and acute intraplaque hemorrhage without administration of contrast agent, provided sufficient spatial resolution and low artifact level (129). Second, non-specific extracellular gadolinium contrast agents are useful in detection and characterization of fibrous cap and neovascularization. Finally, molecular imaging is based on administration of contrast agent tracking specific targets at molecular and cellular level, such as macrophages, thrombus and disease-specific proteins. Molecular imaging agents of different composition such as phospholipid-based molecules and iron oxide conjugates have shown promising results in preclinical testing (130). Expectation for future clinical applications of MRI atherosclerosis imaging is high. Specific targeting of markers of plaque vulnerability at molecular or cellular level (that remain to be determined) would allow identification of patients at high-risk for ischemic events. Following this, MRI as a non-ionizing radiation based modality is promising for monitoring of disease and therapy success by serial studies. Finally, in the research setting, MRI of plaques can be used to study the underlying mechanisms of atherosclerosis and to support the development of new therapeutic strategies (126).

F. The Use of CMR Between Animals and Humans and it’s Application in Specific Settings of Research

By varying the sequences, the contrast and tracking agents, cardiac MR allows the study of many different endpoints, extending from calcium current assessment through to cell tracking and structural characterization. From a translational point of view, a gap obviously exists between the clinical imaging, mainly focused on determining the structure, function, perfusion and viability of the myocardium, and the preclinical studies where imaging is performed at cellular and molecular level (131).

Efforts are being put into the development of clinically compatible imaging techniques, but there is a requirement for additional platforms and there are safety concerns regarding the contrast or targeting agents (132).

Currently, CMR is the only imaging technique allowing high resolution assessment of structure, function, viability and perfusion in a rodent model and is therefore essential for research goals. Technical difficulties arise from the size of the mouse heart (less than one centimeter) and the heart rate of 500bpm. However, with continual technological advancement, CMR is now routinely used in the rodent model (133).

The current applications of CMR are presented in Tables 1-5 with a brief description of the relative positive and negative attributes. Table 1 provides data for myocardial structure and function. Table 2 provides information for myocardial ischemia-reperfusion injury. Table 3 provides a short listing of data for regeneration cardiology. Table 4 outlines information for heart muscle disease and failure, and Table 5 provides data for intervention for arrhythmic or structural heart disease.

G. Summary and Conclusions

CMR has become a powerful tool for use in translational cardiovascular research. Emerging CMR techniques now enable multi-scale characterization of the heart, from changes in global LV structure and function, to changes in tissue composition, mechanics, and perfusion, and exploration of calcium channel function and gene expression. In addition, novel techniques for cell tracking and molecular imaging can be used in both clinical and preclinical research. Many of the techniques described in this chapter can be combined in a single imaging study. A major challenge for the future will involve translation of molecular and cellular imaging to clinical practice.

Acknowledgements

The authors would like to thank Dr Redha Boubertekh and Dr Laura-Ann McGill for their valuable help with images.

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