1 Cardiac MRI Physics

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1 Cardiac MRI Physics

Steven Dymarkowski and Hilde Bosmans

S. Dymarkowski, MD, PhD; H. Bosmans, PhD

Department of Radiology, Gasthuisberg University Hospital, Catholic University of Leuven, Herestraat 49, 3000 Leuven, Belgium

1.1

Basic Principles of Cardiac MRI

1.1.1

Introduction

The basic principles of cardiac MRI are essentially the same as for MRI techniques in other parts of the human body. During an examination, the patient is brought into a high-strength static magnetic field that aligns the spins of the human body (Hendrick 1994). These spins can be excited and subsequently detected with coils. The signal arising from the tissues is influenced by two relaxation times (T1 and T2), proton density, flow and motion, changes in susceptibility, molecular diffusion, magnetization transfer, etc. The timing of the excitation pulses and the successive magnetic field gradients determine the image contrast (Smith and McCarthy 1992).

The purpose of this chapter is to provide insight into the essentials of MR image formation and to discuss specific features of cardiac MRI, rather than to elaborate with a detailed mathematical descrip- tion of every physical phenomenon associated with cardiac MRI.

1.1.2

The Nuclear-Spin Phenomenon, T1 and T2 Relaxation

The asymmetric nucleus of a proton is sensitive to the presence of an external magnetic field. The inter- action of a nucleus with this field can be described via a property that has been called a “spin.” This spin or “magnetic moment” revolves or precesses with a specific frequency, the Larmor frequency, around the magnetic field. The larger the field strength, the higher will be the precessional frequency.

In a large group of spins, such as present in any tissue, the net effect is an alignment of these spins parallel to the long axis of the external magnetic field. A radiofrequency (RF) pulse whose frequency matches

CONTENTS

1.1 Basic Principles of Cardiac MRI 1 1.1.1 Introduction 1

1.1.2 The Nuclear-Spin Phenomenon, T1 and T2 Relaxation 1

1.1.3 Slice Encoding, Phase Encoding, and Frequency Encoding 2

1.1.4 N-space Geometry 4

1.1.5 Spin Echoes and Gradient Echoes 4 1.1.6 Parallel Imaging 6

1.2 Specific Measures in Cardiac MRI Sequences 9 1.2.1 Cardiac Motion 9

1.2.1.1 Synchronization with ECG 9 1.2.1.2 Prospective Triggering and Retrospective Gating 9 1.2.1.3 Single-phase Imaging:

Timing Of Acquisition During Diastasis 10 1.2.1.4 Multiphase Acquisitions 11

1.2.1.5 Real-Time Dynamic MRI 12 1.2.2 Respiratory Motion 13 1.2.2.1 Respiratory Gating 13 1.2.2.2 Breath-Hold Cardiac MRI 13

1.2.2.3 Free-Breathing Navigator-Echo Acquisitions 14 1.3 Contrast Mechanisms in Cardiac MRI 15 1.3.1 T1- and T2-Weighted Techniques 15 1.3.2 Intracardiac Flow and Gradient-Echo Acquisitions 16

1.3.3 Balanced Steady-State Free Precession Sequences 18 1.3.4 Use of Preparation Pulses 19 1.3.4.1 Fat-Suppression Techniques 19 1.3.4.2 Saturation Bands 20

1.3.4.3 Saturation Recovery Pulses 20

1.3.4.4 Dark-Blood or PRESTO Inversion Prepulses 21 1.3.4.5 Contrast-Enhanced Inversion-Recovery MRI of the Myocardium 22

1.4 Specific Protocols 23 1.4.1 Cardiac Morphology 23 1.4.2 Cardiac Function 24 1.4.3 Myocardial Perfusion 25

1.4.4 Myocardial Viability Assessment 26 1.4.5 MR Coronary Angiography 28

1.4.6 MR Angiography of Thoracic Vessels 29 Key Points 30

References 30

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the Larmor frequency can perturb this alignment, resulting in a magnetic moment that is no longer aligned with the field. This magnetic moment of the tissue can now be deconstructed in a vector that is still aligned with the field, the so-called longitudinal magnetization, and a component perpendicular to the field, the transverse magnetization.

After excitation, the spins gradually return to their resting state, i.e., aligned with the field, as this is en- ergetically the most favorable situation. Therefore, their magnetic component along the magnetic field will increase; and the component in the transverse plane will gradually disappear. The first process is described as the T1-relaxation time (Fig. 1.1). A tissue with a long T1 recovers slowly after an RF pulse; the spins of tissues with short T1 align very quickly with the magnetic field. The gradual disappearance of magnetization in the transverse plane after the excitation pulse is called T2 relaxation (Fig. 1.2). Tissues in which the transverse magnetization is rapidly lost are said to have a short T2. These T1- and T2-relaxation processes have a major influence on the image contrast in most MR images (van Geuns et al. 1999).

1.1.3

Slice Encoding, Phase Encoding, and Frequency Encoding

Measurement of T1 and T2 relaxation in itself is not enough to construct an image of the patient.

The coil has in itself no way of determining where in the patient the signals are arising from. This is achieved through spatial encoding. A detailed dis- cussion on how the spatial encoding is performed is beyond the scope of this text, but can be found in any basic MR textbook. The procedure is based on the Larmor equation, namely that the precessional frequency of the spins is proportional to the magnetic field at the location of the spin (Duerk 1999).

In practice, this means that magnetic field gradients are applied over the volume of interest. Let us as- sume the example of reconstructing a transverse slice. To acquire the signal from only this slice, a first magnetic gradient is applied. Along the direction of this gradient (caudal to cranial), the magnetic field, thus also the precessional frequency of the spins, becomes slightly different. As an RF pulse with one

Fig. 1.1. Schematic of T1 relaxation. After the excitation pulse, magnetization is flipped in the transverse plane. During relaxation, longitudinal relaxation (T1) recovers, faster for tissue with short T1 than for tissues with long T1. A second 90°

pulse flips the recovered signal again in the transverse plane, where it can be measured.

Fig. 1.2. Schematic of T2 relaxation. After the excitation pulse, magnetization is flipped in the transverse plane. During relaxation, transverse magnetization (T2) gradually decreases, faster for tissues with short T2 that for tissues with long T2.

A second 90° pulse flips the recovered signal again in the transverse plane, where it can be measured.

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specific RF frequency will excite only the spins with the corresponding precessional frequency, the use of this first magnetic field gradient results in selective excitation of a single transverse slice (slice encoding; Fig. 1.3a).

During the measurement of the signal from this slice, a second magnetic field gradient is applied along another direction (left to right). This is defined as the readout direction. As a consequence, spins of different columns in the excited slice precess with a different frequency, and the measured signal consists of the sum of magnetic moments with different frequencies (frequency encoding; Fig. 1.3b). For example, if this particular signal is measured during 256 successive time points, it is possible to re- distribute the signals afterwards over 256 different columns. The necessary mathematical technique to perform this “column reconstruction” technique is the well-known Fourier transformation.

In order to create differences between the different rows in an image, an encoding scheme has to be used for this direction too. A third magnetic field gradient is applied between RF excitation pulse and

signal readout. It forces the spins of different rows onto a different frequency during a certain time.

The result of this magnetic field gradient will be a different phase for spins from different rows (phase encoding; Fig. 1.3c).

From the above example, it is clear that a basic acquisition scheme is a complex process. For an image with a given resolution of 256², 256 different phase- encoding gradients have to be applied on the same group of spins, and 256 frequency encodings are performed to encode for the horizontal resolution.

A series of successive measurements is therefore necessary. The time in between successive RF excitation pulses is called the repetition time (TR). All the measurements are written in the raw data plane and two-dimensional (2D) Fourier transformation of these data is performed. Because of the common mathematical use of the parameter “N” in Fourier calculations, the raw data space is often called the

“N-space” (Petersson et al. 1993).

The N-space has a central role in MRI. The total amount of data in the N-space determines the image matrix. The direction of filling (bottom-to-top, cen-

Fig. 1.3. a For slice encoding, a magnetic gradient is applied in the Z direction (left). Along the direction of this gradient (caudal to cranial), the magnetic field, thus also the precessional frequency of the spins, becomes slightly different (f1–f6, middle). A radiofrequency (RF) pulse with one specific RF frequency (arrow, right) will excite only the spins with the cor- responding precessional frequency, resulting in selective excitation of a single transverse slice.b For frequency encoding, a second magnetic field gradient is applied along a secondary direction (large arrow). Spins of different columns in the excited slice will precess with a different frequency. To reconstruct the image from this encoded measurement, a Fourier reconstruction needs to be performed. c For phase encoding, a third magnetic field gradient is applied between RF excita- tion pulse and signal readout (large arrow). It forces the spins of different rows onto a different frequency during a certain time. The result of this magnetic field gradient will be a different phase for spins from different rows.

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tric reordered, spiral, etc.) and the sparseness of data sampling (e.g., half-Fourier techniques) have conse- quences on contrast and image resolution (Duerk 1999). Since each line in N-space has experienced different phase-encoding and readout gradients, the information derived from each line is different.

Lines which have experienced the strongest gradients are written near the periphery of the N-space; it is these which make the greatest contribution to the final image resolution. The data samples acquired with the weakest gradients are saved in the center of N-space. They have a major influence on the image contrast.

The different geometries and acquisition schemes using these specific N-space properties is explained in the following section and outlined with detail in relevance to cardiac MRI examinations.

1.1.4

N-space Geometry

The coordinates of N-space are spatial frequencies with units of reciprocal distance, such as millimeter^–1. These spatial frequencies describe how image features change as a function of position in the image (Paschal and Morris 2004). For example: the MRI signal from a large object does not change while passing through its spatial configuration. Otherwise stated: the signal from this object is reflected in low- spatial frequencies. It is only at the border of this object that the transition, e.g., to air or the interface with another object produces a change in signal.

Edges in the images produce high spatial frequencies. The low spatial-frequency information (signal from the object, contrast information) is placed in the center of N-space. The high spatial-frequency information (edges of the object, resolution information) is put outward in the periphery of N-space.

This theory is illustrated in Fig. 1.4. When an image is reconstructed only using the low spatial-frequency information from the center of N-space, the result is a low-resolution image (Fig. 1.4a), while an image reconstructed from only the peripheral N-space lines reveals the borders of the organ with very little contrast (Fig. 1.4b). By gradually adding more peripheral information to the image (Fig. 1.4c), the spatial resolution in the image improves without changing the image contrast. The important elements to re- member from this difficult theory are that the center of the N-space contains image contrast and that the resolution of the image is governed by the extent of peripheral N-space lines (Hennig 1999).

The ways in which N-space can be filled are almost infinite. The classic filling scheme or “trajectory” is sequential bottom-to-top, i.e., the bottom peripheral line is sampled first, continuing in a linear trajectory through the center of N-space toward the top line.

Other schemes include centric-reordered (centrally outward), reversed linear (peripherally inward), radial and spiral. The choice of a particular N-space filling scheme is directed by the information which is to be obtained from the image, e.g., high T2-weighted contrast, first pass of a contrast agent bolus, etc. The different N-space trajectories relevant for cardiac MRI are summarized in Table 1.1.

1.1.5

Spin Echoes and Gradient Echoes

Within the range of sequences used for cardiac MRI, two large families exist. The first commonly used pulse sequence is the spin-echo pulse sequence. In a spin-echo sequence, a 90º pulse is first applied to the spin system. The 90º degree pulse rotates the magnetization into the XY plane. The transverse magnetization begins to dephase. At some point in time after the 90° pulse (usually half of the echo time or TE), a 180° pulse is applied. This pulse rotates the magnetization by 180° about the axis. The 180º pulse causes the magnetization to at least partially rephase and to produce a signal called an echo (Haddad et al. 1995; Fig. 1.5a).

In cardiac MRI, conventional spin-echo techniques suffer from disturbing motion artifacts, even when performed with many signal averages (Winterer et al. 1999). This technique has therefore been largely abandoned. Developments in MRI have gone in the direction of techniques that can be acquired during a single breath-hold. Most frequently, a fast spin-echo technique (FSE/TSE) is used with the TR adjusted to the cardiac frequency. For T1- weighting, the TE is kept short (5–15 ms) and TR is chosen to coincide with one single R–R interval of the patient’s ECG (Stehling et al. 1996; Bogaert et al. 2000). Images are formed by acquiring a number of N-space lines (typically 9–15, named “turbo- factor” or “echo train”) during diastole, in order to minimize motion artifacts. Since the number of N- space lines per excitation pulse determines the total acquisition window per heart cycle, in practice the number of lines has to be adapted inversely proportional to the heart rate of the patient (Fig. 1.5b).

Spin-echo imaging has a major disadvantage.

For maximum signal, the transverse magnetization

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needs to recover to its equilibrium position before TR is repeated. When T1 is long, this can significantly lengthen the imaging sequence.

In a gradient-echo (GE) imaging sequence, first a slice-selective RF pulse is applied to the imaged object. This RF pulse typically produces a rotation angle of between 10º and 90º. A slice-selection gradient is applied with the RF pulse.

A phase-encoding gradient is applied next. A dephasing frequency-encoding gradient is applied at the same time as the phase-encoding gradient so as to cause the spins to be in phase at the center of the acquisition period. This gradient is negative in sign from that of the frequency-encoding gradient turned on during the acquisition of the signal. An echo is produced when the frequency-encoding gra-

Fig. 1.4a-c. Influence of N-space filling on im- age properties. a When only the low spatial- frequency information (center of N-space) is acquired, a low resolu- tion image is obtained. b An image reconstructed from only the peripheral N-space lines reveals the contours of the body and organs with very little contrast. c When the entire N-space is filled, the image has maximal resolution and contrast.

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dient is turned on, because this gradient refocuses the dephasing which occurred from the dephasing gradient (Fig. 1.6).

A period called the echo time (TE) is defined as the time between the start of the RF pulse and the maximum in the signal. The sequence is repeated every TR seconds. The TR period can be as short as tens of milliseconds (Frahm et al. 1990).

1.1.6

Parallel Imaging

Imaging speed is a key factor in most cardiovascular MRI applications of MRI. As a means of enhancing scan speed in MRI, alternative signal acquisition strategies have been developed. These techniques make use of spatial information related to the spatially varying sensitivity of different receiver coils.

Parallel imaging using multiple RF coils has con- siderably enhanced the performance of cardiac MRI. For example, SMASH and SENSE, acronyms for simultaneous acquisition of spatial harmonics and sensitivity encoding, respectively, are two widely used imaging techniques where imaging time has been reduced with a limited sacrifice in signal-to-noise ratio (SNR) (Pruessmann et al. 1999;

Sodickson and Manning 1997; Kyriakos et al.

2000). Both techniques rely on the use of coils with different elements (phased arrays). As the sensitivity of each separate coil element for the geometry of a given object is different, this information can be used to reconstruct information from this sensitivity without actively measuring it, thus saving imaging time. In the SMASH method proposed by Sodickson, the sensitivity profiles of the coils is calculated in one direction. These profiles are then weighted appropriately and combined linearly in

Table 1.1. Different N-space trajectories N-space

trajectory

Variant Mechanism Application in cardiac MRI

Rectilinear Sequential Single line acquired in a separate readout, from top to bottom of N-space

Spin-echo anatomical imaging (older technique)

Turbosequential Multiple lines of N-space per excitation are acquired, each line preceded by a 180°

pulse

TSE or FSE anatomical imaging

Centric- reordered

N-space lines are acquired starting at the center line of N-space and alternating up and down on successive excitations

Intended to acquire image contrast in the begin- ning of the sequence, e.g., first-pass imaging in contrast-enhanced MR angiography

Reversed centric

N-space lines are acquired starting with the outer lines of N-space and working inward from either edge in successive excitations

Intended for high-resolution images with long TE (avoids T2-blurring), e.g., T2-weighted TSE for myocardial edema

Half-Fourier Lines are acquired, usually bottom-to-top up to ±67–75% of N-space. The missing data is replaced by a copy of the bottom part

Faster scanning for which resolution is less important, e.g., fast T2-TSE or HASTE for short-axis determination

Zero filling Usually in a centric reordered sequence, N-space line acquisition is stopped after 50–66% and missing data is replaced by zeros. Equivalent to interpolating a low resolution image to make a higher resolution image

Mainly used in MR angiography to reduce scan time. For example a 144×256 matrix is acquired and reconstructed to 256×256

Radial or propeller

N-space is not traversed line-by-line but rotate around the centerpoint like the spoke of a wheel or a fan blade

Used for sequences requiring high SNR (frequent sampling of the center of N-space) Used in coronary MRA and post contrast T1- weighted 3D imaging (scar imaging) Spiral N-space is usually covered using with

multiple, interleaved spiral trajectories, each after a separate excitations. Known to have complicated reconstruction algorithms

Used for fast acquisitions requiring high spatial resolution. Currently under investigation for coronary MRA

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order to form sinusoidal harmonics which are used to generate the N-space lines that are missing due to undersampling. SMASH has some inflexibility in the choice of imaging planes due its restriction on the placement of receiver coils along one direction, slice geometry, and reduction factor.

The SENSE method proposed by Pruessmann is another parallel-imaging technique which relies on the use of sensitivity profile information in order to reduce image acquisition times in MRI. In SENSE, substantial reductions of scan time are obtained by increasing the distance between the lines in N-space (this reduces the field of view in the image domain).

In doing so, N-space is basically undersampled, resulting in aliasing. Each pixel of an aliased single- coil image reflects the sum of signal components of multiple origins, with each component individually weighted according to local coil sensitivity. The sensitivity-profile information is then used to remove the infolding from the image (Fig. 1.7).

Each method is characterized by a tradeoff in SNR. This is dependent on the data-reduction factor, the configuration and placement of receiver coils, and the accuracy of determination of coil-sensitivity profiles.

Theoretically, the acceleration factor of a parallel- imaging scan is equal to the amount of coils used. In practice however, the degree of scan acceleration is limited by the SNR of the images. In parallel imaging, the SNR is inversely proportional to the square root of the acceleration factor. Furthermore, in image reconstruction from sensitivity-encoded data, noise enhancement occurs in regions where the geo- metric relations of coil sensitivities are not optimal.

Therefore, when parallel imaging is used in cardiac MRI, generally acceleration factors are limited to 2–3.

Parallel imaging is often used to accelerate common breath-hold imaging sequences and to increase the temporal resolution of cine MRI sequences, up to 75 frames/s. Cardiac real-time scanning is one of the greatest achievements of parallel imaging, which will be discussed in Sect. 1.2.1.5 (Weiger et al. 2000). Other applications of parallel imaging in cardiac MRI include decreasing the duration of 3D sequences to fit in a single breath-hold period for assessment of myocardial contrast enhancement (CE-

Fig. 1.6 Schematic of the gradient-echo principle. After the excitation pulse, a phase-encoding gradient and simultane- ously a dephasing frequency-encoding gradient are applied to produce a rephasing signal at the center of the acquisition period (TE).

Fig. 1.5. a Schematic of the spin-echo principle. After an initial 90° pulse, the magnetization is rotated into the XY plane. At TE/2, a 180° pulse is applied, which initiates rephasing. At the echo time (TE), a spin-echo is produced. FID, free induction decay or initial signal after excitation; TR, repetition time. b Schematic of a fast spin-echo (TSE) technique. After the 90°

excitation pulse, multiple 180° pulses (called “echo train” or “turbo factor”) produce a spin-echo signal at TE.

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IR MRI; see Sect. 1.3.4.5) and multislice perfusion imaging.

The most recent advance in this field is to use parallel imaging in the time domain (Tsao et al. 2003).

Due to the intrinsic process of excitation and spatial encoding, in most cardiac MRI sequences, the duration of acquisition of a single time frame is long,

if seen relative to the actual motion of the heart.

Therefore, it can be assumed that there is a certain degree of redundancy in the acquired data. This redundancy is found in the specific spatiotemporal relationship of each individual time frame. Based on this approach, two methods were developed to significantly improve the performance of dynamic

Fig. 1.7 Sensitivity encoding (SENSE) – parallel imaging. Top row: conventional N-space filling (left) with resulting image reconstruction. By reducing the sampling density (increasing the distance between sampling lines in N-space, bottom left), SENSE allows substantial reduction of scan time. Because N-space is undersampled, an aliased image is produced (bottom right). In SENSE reconstruction using phased-array coils, the encoding effect of the different coil sensitivities is taken into account to enable separation of aliased pixels by means of linear algebra.

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imaging, named N-t BLAST (broad-use linear acquisition speed-up technique) and N-t SENSE (sensitivity encoding). In this method, signal correla- tions are learned from a set of training data and the missing data is recovered afterwards. By using this method, it is possible to increase temporal resolu- tions of ungated imaging beyond current capabili- ties during the acquisition stage. It should, however, be remarked that the precise process of temporal parallel imaging is currently not clinically available on most MRI machines.

1.2

Specific Measures in Cardiac MRI Sequences

Compared with other anatomical regions, MRI of the heart faces specific difficulties to overcome, such as cardiac and respiratory motion, flow phenomena, etc. Therefore MRI sequences have been adjusted or are acquired differently for use in cardiac MRI.

1.2.1

Cardiac Motion

1.2.1.1

Synchronization with ECG

The contraction of the heart muscle is a major determining factor in the image quality of cardiac MR images. The application of a conventional MRI protocol would result in unreadable images due to overwhelming motion artifacts (Lanzer et al. 1984).

Using the electrical activity of the heart, MRI data acquisition can be synchronized to the corresponding mechanical events. In this way, cardiac motion can be “frozen”. Different methods and systems are currently available to perform ECG gating in an MRI environment. The user should, however, be aware that the surface electrodes also measure other electrical currents, such as those induced by the magnetic field on the blood flow in the cardiac chambers and vessels. A principal source of artifacts in the ECG due to the magnetic field itself is the magnetohydrodynamic effect. This effect consists of a voltage induced by ions flowing within blood vessels that are exposed to the magnetic field. This voltage artifact is mainly superimposed on the ST segment of the ECG during the ejection of blood in systole. The increase in amplitude of the ST segment

can thus cause a false QRS detection in certain R- wave detection algorithms.

Since accurate peak detecting in the ECG is critically important for good-quality scans, several systems have been designed to provide superior detection. These systems use the spatial information in a vector cardiogram (VCG) to improve R-wave detection in the MRI environment. With a three-dimensional and orthogonal lead system, the electrical activity of the heart is deconstructed in dipolar elements, which allows the R–R interval to be regis- tered as a 3D spatial vector that varies in magnitude and direction throughout the cardiac cycle. Since it has been shown that the average electrical moment of the heart and the average moment of the magnetohydrodynamic artifact manifest significantly different spatial orientations, separation of the QRS loop and the blood flow-induced artifact can be obtained (Fischer et al. 1999).

Other sources of noise in the ECG include that from the RF pulses and the switching of the gradient fields. This has been largely overcome by the introduction of fiber optics in the ECG-detection systems, replacing the formerly used carbon leads. The use of low-power semiconductor optic or laser technology for optimal signal transduction has further improved ECG triggering in the MRI environment.

1.2.1.2

Prospective Triggering and Retrospective Gating There are different ways in which ECG is used to guide acquisition. One technique is to use a preceding R-wave to act as a trigger to acquire information during the following R–R interval or cardiac cycle.

This is called “prospective triggering.” Another strategy is to acquire data continuously and to retrospectively match this to the ECG tracing which is recorded into the memory buffer of the acquisition computer, this is called “retrospective gating.” Both techniques have different applications and will be further elaborated.

Prospective triggering is the more traditional approach: an R-wave serves as a trigger pulse. Instead of acquiring a traditional phase-encoding scheme in which the measurements are performed continuously, by using triggering in cardiac MRI, N-space is segmented into individual lines or groups of lines.

Every group of lines requires another phase-encoding such that they can be saved in the appropriate row of the N-space. An individual group of lines acquired per R–R interval pulse is called a N-space segment. After each excitation pulse, phase-encoding

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tables are incremented and acquisition of a segment is performed (Fig. 1.8a). Depending on the sequence, this can mean that a single segment is acquired for static morphological imaging or a series of segments for multiphase imaging. The key issue to prospective triggering is that the length of the acquisition itself must be shorter that the average R–R interval. If the acquisition length should supersede this period, the next R-wave would be ignored, resulting in un- necessary prolongation of the total acquisition time or breath-hold period, if applicable. For some acquisitions, the changes in effective TR due to changes in R–R interval in case of missed R-wave could also lead to artifacts.

A limitation of prospective triggering is that, due to this constraint, the end of diastole and the early atrial contraction is missed. This can result in poor visualization of physiological phenomena during relaxation and inaccurate readings in flow measurements through the atrioventricular valves. In the clinical setting of diastolic heart failure, restric- tive heart disease, or constrictive pericarditis, these are elements that must certainly not be missed.

Therefore, preference is given to retrospective gating techniques, which acquire the full R-R interval.

In retrospective gating a continuous succession of excitation and data readout is performed and phase- encoding tables are incremented after a fixed time, calculated a priori from an estimated averaged R–R interval. The ECG is recorded during the measurement and, after this measurement, signals are redis-

tributed over the N-space according to their timing over the cardiac cycle. In older sequences, the total acquisition time is usually longer, because in these sequences the actual R–R interval is overestimated to ensure the complete coverage of the heartbeat for each phase-encoding increment. Newer sequences use real-time rescaling of the acquired N-space information to an average R–R interval, and overestimation is no longer necessary (Fig. 1.8b).

1.2.1.3

Single-phase Imaging:

Timing Of Acquisition During Diastasis

In morphological imaging such as T1-weighted imaging of the myocardium and coronary MR angiography, cardiac contraction induces severe motion artifacts related to the length of acquisition if the placement of the acquisition shot is not planned accurately, i.e., during systole or early diastole, when motion of the heart muscle is maximal. Using a time delay, imaging is fixed during the period of diastasis, which is a relative rest period during the mid- to-late diastole. In this period, the heart motion is minimal (Hofman et al. 1998).

The exact timing for diastasis is inversely related to the cardiac frequency of the patient. The length of diastasis decreases as the heart rate goes up. This has implications for artifact-free imaging. In patients with low heart rates, this leaves some margin for error but, in higher heart rates, the length of the

Fig. 1.8a,b Prospective triggering.

In this example of a multiphase acquisition, after each R-wave, a number of N-space lines is acquired.

Note that the acquisition length is shorter than the R-R interval and, at this sampling rate, the end of the cardiac cycle is missed. b Using conventional retrospective gating schemes (R1), the acquisition in- terval supersedes the R-R interval.

This ensures effective acquisition of the complete cardiac cycle, but lengthens the total acquisition duration. In optimized retrospective schemes (R2) real-time rescaling of the acquisition interval to the R-R interval eliminates the need for overestimation.

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acquisition should be decreased to avoid motion artifacts.

There are different ways to calculate the precise timing for diastasis. A first strategy is to calculate the time delay using the empirical Weissler formula. In this formula, trigger delay = [(R–R interval – 350)×0.3]+350.

A much easier method is to perform a cine MRI scan with very high temporal resolution, e.g., 10 ms and to review this in loop mode. Hence the period of diastasis can be assessed as the period of least motion of the ventricle or the coronary artery. The time after the trigger pulse can usually be read in the image information on the scanner and be used as a trigger delay in single-phase imaging sequences.

1.2.1.4

Multiphase Acquisitions

The most common method to acquire dynamic information (e.g., cine MRI or flow measurements) is the single-slice multiphase approach. This approach is typically performed with GE acquisitions, as these can be performed with short TRs in between

successive excitations. These sequences are used to study dynamic phenomena of the heart such as the myocardial contractility and valvular function. The number of phases that can be acquired over the heart cycle are calculated from the acquisition time per phase and the averaged R–R interval. The entire set of images can be loaded into an endless cine loop, providing information on dynamic cardiac processes similarly to other cardiac imaging techniques such as echocardiography, but taking into account the fact that the MR images acquired with this particular scheme are usually not real-time images but are acquired over several heartbeats. Real-time dynamic MRI will be discussed in Sect. 1.2.1.5.

In most acquisition schemes, several N-space lines are acquired per image and per cardiac cycle.

The N-space corresponding to one particular image is then filled with a fixed number of N-lines per heartbeat (Fig. 1.9). This part of N-space is defined as a “segment,” and the corresponding technique is a “segmented acquisition.” Multiphase measurements of successive images can even share N-space data (echo- or view-sharing; Fig. 1.10), which further increases temporal resolution without length-

Fig. 1.9. In a segmented acquisition, a small part of the total image information is acquired after every heartbeat and arranged in N-space. After complete acquisition, the definitive image is obtained by Fourier transformation.

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ening the duration of acquisition. These techniques are usually run during breath-hold.

1.2.1.5

Real-Time Dynamic MRI

With current modern MRI systems and the newest scanning techniques, high temporal and spatial resolution can be achieved in cine MRI sequences. Using a SSFP protocol (Sect. 1.3.3) and parallel imaging, a complete set of cine MR images can be obtained in 5–10 breath-hold periods of 10 s. Nevertheless, this procedure can be complicated by arrhythmias and patient difficulty with long or multiple breath-holds, especially in a group with cardiac illness.

For cine MRI, sequences have been adapted allowing for imaging in real time. In a conventional N-space sampling strategy, each image is composed of N-space segments acquired over several R–R in- tervals during cardiac triggering or gating. In real- time sequences, N-space is generally filled in a single shot. The imaging matrix is reduced. e.g., 128², no cardiac triggering is used, and parallel imaging is used to increase the data sampling rate. A single image can thus be acquired in less than 100 ms. Since the acquisition duration is short, breath-holding is optional, so it can be used even in critically ill pa-

tients. Studies have shown this technique to produce comparable results with standard multiple breath- hold examinations (Fig. 1.11; Spuentrup et al. 2003;

Hori et al. 2003).

A further evolution of this technique tackles another potential source of errors both in segmented breath-hold and free-breathing scans, namely slice misregistration between multiple successive slices.

Depending on the imaging plane, diaphragmatic shifts between breath-holds of unequal depth or during free breathing can cause cardiac shifts of up to 1 cm. This could result in over- or underestima- tion of cardiac volumes during quantification.

The abovementioned real-time scans can be combined with retrospective gating to produce a real-time, retrospectively gated, multislice scan with retrospective reordering of the acquired real-time images into the R–R interval over which they were acquired. A further advantage over nontriggered scans is that these images can be accepted by postprocessing software as triggered cardiac studies and can be quantitatively analyzed.

Similar evolutions exist for myocardial perfusion sequences, which visualize the first pass of contrast agents in the myocardium (see also Chap. 7). High spatial resolution is not a basic requirement for perfusion studies. The major prerequisites for perfu-

Fig. 1.10. View sharing. Each individual cine MR image is formed from different N-space segments. Virtual images are formed from the half of each preceding and following segment through N-space interpolation.

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sion acquisitions are a short acquisition time and an adequate T1 weighting. In practice, single-shot T1- weighted gradient-echo sequences are used. These sequences use a saturation-recovery prepulse, a 90º pulse which guarantees the nulling of the entire signal prior to the effective acquisition, allowing clear visualization of the first-pass effects.

1.2.2

Respiratory Motion

Respiration and the associated motion of the organs in the thorax and the abdomen are well-known prob- lems in MRI. This kind of motion superimposes on the signal acquisition during cardiac contraction and is of approximately equal importance for image quality.

From early experience with MRI, increasing the number of signal averages decreased motion artifacts in the images. However, as the total acquisition time is proportional to this number, this approach was very time-consuming. In addition, even for properly adjusted T1-weighted spin-echo techniques, good results could never be guaranteed. Since in the present era most imaging techniques used in cardiac MRI use accelerated acquisition schemes (fast spin-echo, turbo-gradient echo) other strategies are maintained to keep images clear of respiratory artifacts (Pettigrew et al. 1999).

1.2.2.1

Respiratory Gating

Similar to cardiac triggering, acquisitions can be triggered by the respiratory cycle. By use of a respiratory belt, excitation and signal readout is guided to the period of end-expiration. In cardiac MRI,

this technique presumes that cardiac and respiratory gating are combined. The disadvantage of this approach is that acquisitions become very long. As one respiratory cycle can easily take 5 s or longer, the minimal effective TR for an acquisition would be 5 s.

In practice, this technique is rarely used.

1.2.2.2

Breath-Hold Cardiac MRI

The developments in MRI tending toward fast-acquisition schemes make imaging during a single breath-hold possible. Generally speaking, a ran- dom patient can hold his or her breath for about 10–15 s. The goal is to use an acquisition that can be performed in this time frame. Clear patient instructions before the examination can aid in improving the efficiency of this approach. It can also be useful to provide patients short of breath with an oxygen mask or nasal cannula during the examination to alleviate the often frequent breath- holds.

For morphological imaging in breath-hold, usually a FSE or TSE sequence is used, in which the number of lines per heartbeat can be tailored to coincide with the period of diastasis in diastole, as described in Sect. 1.2.3. If T2-weighting is required, one should consider that, in view of the long echo times, there is a linear relationship between the length of the acquisition shot and the amount of blurring in the image due to dephasing. A compromise should be found between the length of breath-holding and the length of the acquisition. If necessary, parallel- imaging technology can be used to shorten the total acquisition duration. Similar acquisition schemes in GE protocols allow for cine MRI and flow measurements to be captured in a single slice per breath- hold.

Fig. 1.11. Comparison of retrospective cine MRI (left) and real-time un- gated cine MRI (right).

The ungated image has lower spatial resolution, since priority during the acquisition is given to temporal resolution.

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1.2.2.3

Free-Breathing Navigator-Echo Acquisitions

In longer imaging sequences such as high-resolution flow measurements and especially in coronary MR angiography, abovementioned strategies have little chance of success. The long acquisitions need a different approach to filter the respiratory motion out of the acquisition. This is performed by navigator- gated scans. This technique will be briefly described here, since it is elaborated in Chap. 14. A navigator is usually a 2D RF pulse, representing a single signal readout of a sagittal slice, positioned over the right diaphragmatic dome. It is positioned there in order not to pick up any signal from the heart. The signal readout of this RF pulse allows determination of the position of the diaphragm, since there is high contrast between the liver and the lungs. During the acquisition of the imaging sequence, this signal or position

is read out every heartbeat, thus providing a real-time measurement of the diaphragmatic excursion. By use of a threshold value within which range to accept data, the MRI system can be directed only to include information into the N-space of the given protocol, for instance at end expiration. All other data are rejected (Fig. 1.12). The length of the total acquisition depends on the efficiency of the navigator and thus on the regularity of the breathing pattern of the patient. If respiratory drift occurs, for instance when the patient falls asleep, or there is bulk motion, the efficiency is limited. It is therefore important to carefully instruct the patient to breathe regularly and consistently.

There are many variations in navigator gating.

Some tracking algorithms exist to compensate for drifts in respiratory excursion, and many investiga- tors have works with navigator beams positioned on the heart itself. Additional information about coronary MR angiography can be read in Chap. 14.

Fig. 1.12 Principle of navigator echo acquisitions. A naviga- tor pencil beam is placed on the right diaphragmatic dome as shown on the top part. The navigator registers the motion of the diaphragm during respiration and determines end-expiration as the top of the excursion curve. All data points of the acquired sequence within a certain acceptance window (indicated by the blue lines) are kept for image formation, while other points are rejected.

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1.3

Contrast Mechanisms in Cardiac MRI

1.3.1

T1- and T2-Weighted Techniques

Static images in cardiac MRI are often used for anatomical evaluation. These images require high image contrast and a sufficient spatial resolution. The required weighting is in practice application-dependent. For instance, the visualization of myocardial edema is optimal with a T2-weighted acquisition, whereas other indications such as evaluation of the right ventricular (RV) wall, LV wall thickness, and pericardial thickening benefit from T1-weighted measurements. The visualization of enhancing structures after an intravenous injection of contrast agents (infarction, mass, fibrosis) also requires a T1-weighted approach (Fig. 1.13a,b).

Conventional spin-echo techniques suffer from disturbing motion artifacts, even when performed with many signal averages and a long TE. This technique has therefore been largely abandoned.

Developments in MRI have gone in the direction of techniques that can be acquired during a single breath-hold. Most frequently, a fast spin-echo technique (FSE/TSE) is used with the TR adjusted to the cardiac frequency. For T1-weighting, the TE is kept short (5–15 ms) and TR is chosen to coincide with one single R–R interval of the patient’s ECG. Images

are formed by acquiring a number of N-space lines (typically 9–15) every diastole, in order to minimize motion artifacts. Since the number of N-space lines per excitation pulse determines the total acquisition window per heart cycle, in practice the number of lines has to be adapted inversely proportionate to the heart rate of the patient.

GE acquisitions can be used as an alternative to T1- weighted TSE acquisitions. A GE or turbo-FLASH (fast low-angle shot technique) acquisition can be used as a single-shot or multishot technique with TR/TE minimal and an inversion-recovery (IR) preparation pulse to improve the image contrast (Frahm et al. 1990).

T2-weighted acquisitions can also be performed with fast acquisition schemes. Both fast spin-echo (TSE/FSE) and echoplanar imaging (EPI)-based techniques have been explored. The turbo-factor or number of N-space lines that can be accepted for T2- weighted imaging is subject to the same type of compromise as T1-weighted acquisitions. For T2-weighted TSE, the turbo factor in TSE imaging is usually higher (21–33) and TR is longer (2 or 3 heartbeats), which is beneficial for T2-weighted contrast. These measurements are often combined with fat-suppression techniques to increase image contrast (Stehling et al.

1996) To increase the contrast in T1- and T2-weighted images, these sequences can be fitted with a double inversion or “black-blood” pulse to increase image contrast between cardiac structures and the blood.

This functionality is discussed in Sect. 1.3.4.4.

Fig. 1.13 T1- and T2-weighted techniques. Left: T2-weighted TSE image with fat suppression acquired in a patient with acute myocardial infarction. The tissue edema due to the infarction is expressed as elevated signal intensity in the anterior wall.

Right: T1-weighted gradient-echo acquisition after injection of contrast in a patient with asymmetric septal hypertrophy.

Disarray of myocardial fibers and increase in interstitial fibrosis is shown as a contrast-enhancing wall segment.

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Postcontrast T1-weighted MRI, for instance in the setting of ischemic heart disease, is used to distinguish between myocardial infarction and normal myocardium. Different authors have shown that visualization of contrast enhancement can be best performed using GRE acquisitions. Newer MR sequences, such as inversion-recovery fast GE (e.g., IR- FLASH, IR-FFE, CE-IR MRI) sequences can be used to obtain T1-weighting, while the inversion-recovery prepulse (inversion time, TI, variable between 200 and 300 ms, depending on contrast dose and time after injection) suppresses the normal myocardium and thus improves the contrast difference between normal and pathological myocardium (Fig. 1.14) (Sandstede 2003). This technique will be further elaborated in Chap. 8, Ischemic Heart Disease.

1.3.2

Intracardiac Flow and Gradient-Echo Acquisitions

Most conventional MRI sequences presume that the spins in the tissues do not move during the procedure.

For flowing spins, such as those in the ventricular cavity, the great vessels, or in the coronary arter- ies, this assumption is not valid. Spins moving along the direction of a magnetic field gradient acquire a phase angle that depends on the magnetic field and on their flow. A paradoxical signal enhancement of flowing blood is observed in GE acquisitions whenever the spins experience only a restricted number of RF excitation pulses and then disappear from the imaged volume. Hence, the continuous refreshment or inflow of the spins ensures a maximal alignment of spins with the external field prior to the excitation

pulses. Compared with the partially relaxed signals in the stationary tissues, subjected to a long series of RF pulses, the signal in the vessel or cardiac cham- ber can therefore be very strong. Apart from being the basis for time-of-flight (TOF) MR angiography acquisitions, this principle is used for GE cine MRI sequences (Fig. 1.15; Semelka et al. 1990). Usually single-slice, segmented multiphase GRE protocols or hybrid variations (GRE-EPI) with small flip angles are used for this purpose. The myocardium can be delineated due to hyperintense signals in the cavity, and these sequences possess a sufficiently high temporal resolution to acquire 15–20 frames per cardiac cycle. Sharing of the measured raw data (echo- or view-sharing) increases the temporal resolution further (Bogaert et al. 1995).

There have been a number of alternative acquisition schemes proposed over the last 5–10 years, such as multishot echo planar MRI (EPI), IR single-shot turbo-FLASH and biphasic spin-echo sequences as compared to the well-validated conventional GRE technique. Although these techniques all provided accurate measurements of global LV function and mass in a time-efficient manner, they are now largely abandoned for balanced steady-state free precession (b- SSFP) sequences, which are discussed in Sect. 1.3.3.

The flow-induced phase effects caused by extra bipolar pulses in GRE sequences can be exploited to study intracardiac or valvular flow. Due to a bipolar pulse, flowing spins acquire a phase angle that is, in a first approach, proportional to the amplitude of the gradients, the time between the onset of the first and the second pulse, and, most importantly, the velocity of the spins (whenever their motion can be described by a constant velocity flow during the application of the gradients). Under well-controlled

Fig. 1.14 Contrast-enhanced inversion-recovery (CE-IR MRI) technique. Since longitudinal relaxation is tissue- specific, a 180° (inversion) prepulse can be used to en- hance the contrast between normal (dark) and scarred (bright) myocardium as shown in a patient with an old inferior myocardial in- farctation. Midventricular short-axis view, obtained 20 minutes following injection of 0.2 mmol Gd-DTPA/kg body weight, using an inversion time of 270 ms.

Mz

Mxy

D 180

TI or inversion time null point

(tissue specific)

scarred myocardium

normal myocardium

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Fig. 1.15 Gradient-echo cine MR image in cardiac short-axis. There is moderate blood-to- myocardium contrast, mainly due to intracavi- tary flow.

c

Fig. 1.16a–c Gradient-echo phase-contrast flow measurement of the aortic valve. a Phase-encoded image. Flow in cranial direction is displayed as white pixels, while flow in the caudal direction as black pixels. Stationary tissues appear as intermediate gray-scale image points. b corresponding magnitude image. c flow velocity curve obtained by measuring the flow profile over time.

conditions, the velocity of the spins can thus be measured from the phase angles (Higgins et al. 1991).

A typical flow-encoded phase image is shown in Fig. 1.16a. Measuring the phase angle over the cardiac cycle permits the calculation of the mean flow

over, for example, the descending aorta, as shown in Fig. 1.16b. The range of velocities that can be measured is determined by the gradients and generally described by the strength of the VElocity-ENCoding bipolar pulse (VENC).

a b

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1.3.3

Balanced Steady-State Free Precession Sequences

Since the introduction of powerful gradient MRI systems into clinical practice, great interest has re- emerged for b-SSFP techniques. These sequences have many monikers, such as TrueFISP (Siemens), balanced FFE (Philips), and FIESTA (GE). As very shrewdly described by Scheffler et al., many authors unify these sequences under the name “steady-state free precession techniques” (SSFP) and emphasize that the unique feature of this sequence is the acquisition of magnetization during steady state.

Scheffler accurately remarks that is very confusing and wrong, since all rapid GE sequences are SSFP sequences (Scheffler and Lehnhardt 2003). The remarkable features and contrast mechanisms of the b-SSFP sequence do not originate from the steady state but from the unique gradient-switching pattern through which it generates contrast.

As previously described, a GE is made up of a number of excitation pulses, separated by a given time interval (TR). The acquisition and the spatial encoding of the GE is performed between consecutive excitation pulses by means of switched gradient pulses along read, phase, and slice direction. The amount of signal dephasing within TR depends on the spatial position and on the gradient strength. In addition the spatial specific dephasing, T1 and T2 relaxation unavoidably occurs during TR, resulting in further signal decay. The next excitation pulse is superimposed on the altered magnetization, and the process of excitation and relaxation is repeated every TR. After several TR periods – providing this is a constant period – a steady state of the magnetization is established (Fig. 1.17). This situation is called

“steady-state free precession” (SSFP). All types of fast GE sequences are SSFP techniques. The difference between the various types of GE sequences is a different gradient-switching pattern between the consecutive excitation pulses. Different gradient-switching pat- terns produce different dephasings within TR, which finally result in different types of steady states and, more importantly, in different image contrasts.

b-SSFP is a special type of SSFP sequence where the gradient-induced dephasing within TR is ex- actly zero. In other words, within TR each applied gradient pulse is compensated by a gradient pulse with opposite polarity, which ensures maximum recovery of transverse magnetization. The overall magnetization consists of a single vector that is not spatially dephased, as for non-balanced SSFP

sequences. Absence of this linear dephasing of the magnetization as seen in conventional excitation schemes results in much higher signal gain in the image (Fig. 1.18).

Besides a favorable SNR, the most important feature of an imaging sequence is its contrast. While the classic spin-echo sequence shows either a T1- or a T2-weighted contrast, rapid GE sequences, includ- ing b-SSFP, exhibit a relatively complicated contrast that is composed of T1 and T2 contributions (Hennig et al. 2002). In cardiac MRI, due to the transient signal of inflowing blood and the steady- state signal of muscle (which is low due to a low T2/

T1 ratio) can be observed, which makes b-SSFP very useful for cardiac imaging. The high contrast of b- SSFP is produced both by the different ratio of T2 and T1 of blood and myocardium and by inflow effects. A second advantage of b-SSFP is its high SNR even for very short TR. A GRE sequence with comparable TR and optimized flip angle exhibits a much lower SNR and contrast-to-noise ratio.

The new b-SSFP technique is presently the pre- ferred cine MRI technique to study ventricular function, since it yields a better contrast between blood and myocardium, thus providing better visualization of small anatomical structures such as endo- cardial trabeculations, papillary muscles, and valve leaflets (Thiele et al. 2001). Combined with parallel imaging (SMASH, SENSE), the temporal resolution can be increased by a factor of 2–4 (i.e., temporal resolution <10 ms), albeit at the expense of SNR.

Fig. 1.17 Schematic of a balanced steady-state free preces- sion (b-SSFP) sequence. In this acquisition scheme, all three gradient axes are refocused or balanced. If compared with non-balanced SSFP (Fig. 1.6), in b-SSFP the readout gradient is perfectly symmetrical (arrows), resulting in completely rephased magnetization after one TR period.

Gradient

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b-SSFP requires short TRs (3–6 ms) to minimize artifact behavior. This is the main reason why b- SSFP has only become feasible (and popular) during the past 3–5 years. With the design of very fast gradient amplifiers, it was possible to reduce the time to switch all relevant gradients needed for echo encoding to a few milliseconds. The b-SSFP also requires relatively high flip angles, between 50 and 80°, to generate the highest possible signal. This can easily exceed the SAR limits, which may limit its use in ultrahigh field systems, such as recently introduced 3 T MRI units.

1.3.4

Use of Preparation Pulses

Preparation pulses create a preexisting condition in the spin population prior to the applications of RF pulses destined for signal readout. Typical ex- amples are fat-suppression pulses, inversion pulses, and saturation pulses. Other types of preparation pulses have applications in other parts of the body, such as magnetization transfer pulses, vessel tracking pulses, spin-labeling pulses, but are beyond the scope of this chapter.

1.3.4.1

Fat-Suppression Techniques

Fat suppression in cardiac MR images may be desir- able for a variety of reasons. An examination aimed at the diagnosis of right ventricular pathology may require fat-suppressed TSE images to distinguish the fine anterior wall from the pericardial fat. Fat suppression is commonly used in coronary MR an-

giography and in contrast-enhanced sequences used for myocardial viability assessment.

A robust way to perform fat suppression is to use sequences with an IR prepulse. These sequences are called STIR sequences, which is an acronym for short-tau inversion recovery. This sequence starts with a 180º RF pulse that inverts the magnetization (Simonetti et al. 1996). After this pulse, the spin population relaxes: it realigns its magnetization with the external magnetic field. This relaxation behavior is, by definition, described by the T1 relaxation time. A time constant “TI” is associated with the use of this pulse: it defines the time in between the IR pulse and the actual measuring process. This sequence parameter could be set to optimize the T1- weighted contrast. Fat suppression is achieved with a TI of about 120 ms. At this TI, the signal of the fat crosses the zero point. STIR is most commonly explained with reference to an inversion-recovery spin-echo (IRSE) sequence, but the approach is also applicable to TSE, GRE, and EPI acquisitions.

IR prepulses can also be added to a T2-weighted sequence, for instance to visualize myocardial edema in the setting of an acute myocardial infarction (Dymarkowski et al. 2002). The combina- tion of T2-weighting, usually in the form of a TSE sequence and fat suppression causes high contrast between the edematous myocardium and the surrounding tissues. Beware, however, that overall SNR decreases as TE is increased. For adequate SNR, an echo time of 60ms can be chosen (Fig. 1.19).

In spectral fat suppression, or SPIR, a spectrally selective “fat-pulse” is used to flip fat spins and, after the time interval that lets the longitudinal magnetization of fat reach zero, the excitation pulse is applied which causes the signal of the water spins to contribute to most of the signal. The best fat sup-

Fig. 1.18 Balanced steady-state free precession (b-SSFP) cine MR image in cardiac short- axis. In comparison with standard gradient-echo cine MRI, the blood-to- myocardium contrast is higher. This technique allows better delineation of intracardiac structures such as papillary muscles and trabeculations.

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pression is achieved with a 180 spectral inversion pulse and a time delay equal to that used in normal STIR imaging (Fig. 1.20).

1.3.4.2

Saturation Bands

Saturation bands are normal 90º RF excitation pulses that are applied on tissues prior to the measurement, thereby avoiding interference of their signal with the images. This is an efficient way to suppress image artifacts originating from a traceable source that does not have to be visualized in the image. A commonly used application of saturation bands is slabs positioned over the chest wall to avoid high signal from subcutaneous fat in T1-weighted imaging or to avoid motion-related artifacts in coronary MR angiography (Fig. 1.21). Another very interesting application of these saturation bands is the selective saturation of parts of the myocardium, as first de-

scribed by Zerhouni et al. The saturated tissues can be used as noninvasive markers or “tags” to study the myocardial deformation and motion during the heart cycle (Zerhouni et al. 1988; see Chap. 6).

1.3.4.3

Saturation Recovery Pulses

The saturation recovery (SR) sequences are only rarely used for imaging. Their primary use at this time is as a technique to measure T1 times more quickly than an IR pulse sequence. SR sequences consist of multiple 90º RF pulses at relatively short TRs. Longitudinal magnetization after the first 90º RF pulse is dephased by a spoiling gradient (in this case with the slice select gradient). Longitudinal magnetization that develops during the TR period after the dephasing gradient is rotated into the transverse plane by another 90º pulse. A GE is acquired immediately after this. The signal will reflect T1

Fig. 1.19 Use of fat sup- pression in T2-weighted MRI. Left: T2-weighted TSE image in the cardiac short-axis in a 64-year- old patient with acute myocardial infarction.

In this image without fat suppression, the tissue edema is only margin- ally visible (arrow).

Right: corresponding fat-suppressed image shows increased contrast between the edematous and normal myocardium (arrowheads).

Fig. 1.20 In CE-IR-MRI, spectral fat suppression is used to suppress the signal of chest wall (arrow) and pericardial (arrowhead) tissue to in- crease contrast between the myocardium and surrounding fat.