RADIOLOGYFig. 1.55. The time-of-flight effect in 3D magnetic resonance angiography (3D ToF-MRA). The contrast is given by the amount of blood that is flowing into the excited slab, replacing the saturated fluid. For a thick slab and slow flow, there is the potential of saturation towards the distal portion of the slab. This effect is reduced by using an asymmetric radiofrequency (RF) slab profile, also called TONE (tilted optimized non-saturating excitation). In that case, the excitation angle at the entry point of the vessel is lower than at the exit port. This technique depends on the direction of flow.

Fig. 1.56. The maximum intensity projection (MIP) projects the maximum intensity found in a stack of images to the pixel intensity on the screen, depending on the view, the perspective, that the user defines. In this example, a transverse 3D ToF-MRA was performed to study the aneurysm in the right vertebral artery. A coronal MIP was applied to these transverse stack of images.

Areas and for the smaller vessels, there is always the possibility of assessing the native images. This 3D ToF-MRA technique is still used for imaging the intracranial celebral circulation.

A method to improve the contrast between the vessel lumen and the surrounding brain parenchyma is the application of a MTS pulse as discussed in Sect. 1.1.1.4. The stationary brain parenchyma contains macromolecules which are ‘invisible’ due to a short T2, bounded water molecules that can be saturated via an off-resonance RF pulse. Via magnetization transfer, this leads to an increased background suppression. The progressing saturation of the flowing blood from the entry point towards the exit point is usually compensated with a linear flip-angle change in the direction of the expected flow: the RF pulse that excites the whole 3D volume is designed to provide a low flip-angle excitation at one side of the 3D slab, the entry point of the vessel, with an increase of flip angle towards the other side of the slab. The aim of this procedure is to achieve a homogeneous signal contribution throughout the 3D volume. Such a technique is also called tilted optimized non-saturating excitation (TONE).

1.2.2

2D Time-of-Flight Angiography

Shortly after the introduction of the 3D MRA method, the 2D technique was presented as being the better method for visualizing the extracranial celebral circulation and the peripheral arteries. The main source of contrast in 3D ToF-MRA is the inflow of unsaturated blood into the volume. The drawback with the 3D ToF-MRA technique is the progressive saturation of blood that travels through the volume. The idea of the 2D ToF-MRA technique is to make every slice an entry slice as illustrated in Fig. 1.57, improving the contrast dramatically compared with the 3D ToF-MRA technique. One disadvantage of the 2D-ToF technique is the prolonged measurement time or the limited coverage of the sequentially applied slices. These slices have to

be acquired with a significant overlap in order to compensate for the signal distortions at the edges of the slice. Otherwise, a typical ‘staircase’ pattern is to be observed in the maximum intensity projection (MIP). Another severer limitation of the 2D ToF-MRA technique is the rapid saturation of blood that is coincidentally flowing within the slice or reentering the slice in the case of a vascular loop, mimicking the lack of flow as found after stenoses or occlusions. The evaluation of these data sets are again performed with a MIP technique as discussed in the 3D ToF-MRA section. The 2D ToF-MRA methods are being replaced by the contrast-enchanced MRA techniques, since ceMRA protocols are less complex and much faster, and the resulting images demonstrate fewer artifacts.

Fig. 1.57. The 2D time-of-flight magnetic resonance angiography technique (2D ToF-MRA) avoids the saturation problem found in 3D applications by acquiring thin 2D slices sequentially. Each slice is therefore an entry slice. The-slices are usually acquired with a significant overlap in order to compensate for signal variations at the edges of the slice caused by non-ideal slice profiles. RF, (radio frequency).

1.2.3

3D PC Angiography

MRA techniques based on a velocity-dependent phase shift of the transverse magnetization, the phase-contrast (PC-MRA), use a slightly detuned GMR arrangement. The transverse magnetization within voxels containing flowing blood will have a different phase position than the macroscopic magnetization within the voxel of adjacent stationary tissue. It is usually the ‘different vector’ that is utilized for the visualization of the vascular tree (Fig. 1.58). The reference phase is usually acquired with a preceding scan with perfect GMR. In order to obtain all three possible velocity components, the ‘detuned’ GMR measurements have to be applied to all three possible orientations. The advantages of this technique are the perfect background suppression and the adjustable sensitivity to slow velocities. The major disadvantages of this technique are the relatively long measurement times and the potential selection of an improper flow sensitivity. If the selected sensitivity is too high, corresponding to an underestimated velocity range, the contrast will be poor. If the selected sensitivity is too low, corresponding to an overestimated velocity range, the contrast will be poor. In addition, this technique is sensitive to higher-order motion and apparently not as robust as the ToF techniques. The PC-MRA methods are being replaced by the contrast-enchanced MRA techniques, since ceMRA protocols are much faster, and the results are more consistent.

1.2.4

2D PC Angiography

The 2D PC-MRA technique takes advantage of the same phenomenon. Rather than acquiring multiple 3D partitions, however, only one thick slice is selected. The advantage of

such an approach is the relatively short measurement time compared with the 3D method. Another advantage is that the image provided is already the angiogram. There is no need and no possibility of further post-processing, such as the presentation of a different view angle. The latter is a disadvantage of the 2D PC-MRA technique. Another disadvantage is a poorer SNR compared with a 3D acquisition. The fact that the voxel in 2D PC-MRA is highly anisotropic also leads to a potential pitfall of adding the contribution of various vessels flowing through the same voxel. 2D PC angiography techniques are usually taken to achieve a fast angiographic localizer.

Fig. 1.58. Principles and example of a phase-contrast magnetic resonance angiography (PC-MRA) technique. The left graph and image present a so-called PC-MRA of the lower extremity. The length of the vector between the stationary reference and the signal from the moving blood is translated to a pixel intensity within the image. The right graph and image demonstrate a so-called phase map (or phase-different image). The basic sequence structure is identical, but inthis case, the phase difference is converted to a signal intensity within the image. The gray scale in this case corresponds directly to a measured velocity. The transverse cut through the aortic arch at the level of the pulmonary artery is shown.

1.2.5

Contrast-Enchanced Magnetic Resonance Angiography

With the new progress in hardware development and the ability to acquire a 3D MRA within a breath-hold-or the passing of a contrast bolus-the so-called contrast-enchanced MRA techniques have shown a dramatic improvement, especially regarding the abdominal and peripheral vasculature. The technique itself is rather trivial. A GRE technique is applied with the shortest suitable TE selected , the shortest possible TR, and a moderate excitation angle. The aim is to provide an image of a passing contrast bolus. Figure 1.59 demontrates an example of a so-called time resolved ceMRA, where the measurement of the transit time is omitted and replaced by multiple sequential acquisitions of the same region. The T1 shortening of the blood as a consequence of administering a paramagnetic contrast agent will allow the imaging of the vascular tree with no saturation effects. There is no entry plane and no concern about saturation. The challenging part for the user is the timing between the injection of the bolus and the start of the breath-hold acquisition, with the goal of acquiring the low k-space frequencies at the time the bolus passes through the region of interest. The timing of the measurement depends on the transit time of the bolus to the region of interest, the phase-encoding loop structure of the imaging sequence, and the duration of the injection time. The transit time, the time between intravenous injection of the contrast bolus and the appearance within the region on interest, can be evaluated prior to the ceMRA protocol by using a small test bolus and turbo FLASH technique through the region of interest with an update of 1 images/s. As alternative is usually offered in the form of an interleaved sequence. The user can start a bolus tracking fast 2D technique (e.g., turboFLASH), inject the contrast bolus, and semi-automatically can switch to the (3D) ceMRA technique as

soon as the arrival of the bolus is noted within the region of interest (e.g., CARE bolus technique). To cover large regions as required in the ceMRA evaluation of the peripheral vasculature, protocols are offered that include automatic table feed to cover several vascular stations utilizing one single bolus injection (panoramic table MRA). A more detailed protocol optimization for large vessels and peripheral vessels is discussed in Chapter 13.

Fig. 1.59. ceMRA Principle. Contrast strongly depends on matching the arrival of the bolus within the region of interest and placing the k-space acquisition accordingly. Time-resolved imaging of the extracranial cerebral circulation will provide the native scan, the arterial phase, and images of the veno is phase of the passing contrast media.

1.2.6

Flow Quantification

Since the phase change of a de-tuned GMR is proportional to the velocity of the flowing blood, this phase difference can be used for quantification. Although rarely applied in the clinical routine, it has some future potential in quantifying the flow through a dialysis shunt, investigating valvular insufficiencies, grading of shunts in congenital malformations of the heart, providing information on the extent to which flow in false lumen supplies vital organs in aortic dissection, and investigating patent CSF channels in patients with hydrocephalus.

1.3

Techniques in Cardiac Imaging

The time-consuming acquisition of multiple Fourier lines usually takes too long to capture the motion of the beating heart. Exceptions are echo planar imaging techniques and recently developed ‘real-time’ trueFISP sequences. For all other imaging techniques, image acquisition is triggered or gated with physiological signals from ECG electrodes or a pulse sensor. The advantage of the time-of-flight effect of inflowing blood, creating a hyperintense signal appreciated in MRA, turns into a disadvantage for certain cardiac applications and is compensated with a so-called ‘dark blood’ magnetization preparation scheme.

1.3.1

ECG Gating-Prospective Triggering and Retrospective Cardiac Gating

In order to get a ‘frozen’ image of the beating heart, the Fourier lines for the slices to be imaged have to be taken at the same point in time within the cardiac cycle. The starting point for a multi-slice measurement of T1-weighted images acquired with conventional spin-echo sequences is estimated prospectively based on the advent of the ECG signal as illustrated in Fig.1.60. As a practical hint, the acquisition of a stack of slices should be moved towards to the end of the cardiac cycle in order to minimize motion artifacts. Conventional spin-echo sequences

are usually optimized for cardiac imaging using an additional pair of dephasing gradients in order to diphase the signal from the moving blood. Parallel saturation band around the stack of slices are used to minimize the inflow effect of unsaturated blood.

With the introduction of fast low-angle shot imaging (FLASH) and other gradient-echo techniques, it became possible to reduce the TR to e.g., 40ms and lower, allowing the acquisition of time-resolved images of the beating heart. Figure 1.61 illustrates the prospective triggering previously described as well as the retrospective cardiac gating approach. For the latter, Fourier lines are measured for the same phase-encoding step with the selected repetition time of the sequence and are stored together with a time stamp of the last ECG event. After completion of one or two cardiac cycles, the phase-encoding amplitude is advanced to measure the next Fourier line. Data are later normalized and resorted, and the user can select the temporal resolution, that is, the number of images that he would like to have calculated for one cardiac gating is that the cardiac cycle is covered completely, without any gap in time, whereas for prospective triggering there is a gap between the last measurement and the beginning of the next measurement with the next ECG event. The disadvantage of retrospective cardiac gating is the sensitivity to extrasystolic events and arrhythmic heart beats. The measurement may become invalid if such an event occurs when the center of k-space is being acquired.

Fig. 1.60. Prospective ECG-triggered multi-slice acquisition. For simplicity, the k-space consists of only three Fourier lines (three different gray-shaded boxes) in this illustration. The usual matrix size reguires 128 Fourier lines. One Fourier line is measured per heart beat. Multiple slices can be measured within one heart beat. Measurement should take place in end-diastole to minimize motion artifacts.

Fig. 1.61. Prospective ECG-triggered single-slice acquisition and retrospective cardiac gating. In this illustration, the k-space consists of only three Fourier lines (three different gray-shaded boxes). The usual matrix sizes requires 128 Fourier lines. In prospective triggering, one Fourier line is measured per heart beat per cardiac phase. In retrospective cardiac gating, the same Fourier line is measured continuously well beyond the time of a cardiac cycle. The Fourier lines are later normalized and resorted. The temporal resolution is given by the number of images per cardiac cycle as selected by the user. Images are reconstructed based on interpolated and weighted Fourier lines measured within the given time segment.

1.3.2

Segmentation and Echo Sharing

Unfortunately, the heart is not only beating, its position also depends on the breathing cycle. In conventional imaging, two to three averages are used to smooth the artifacts based on respiratory displacement. As a consequence, the measurement time is prolonged, and the outline of the myocardial border becomes soft. In order to reduce the measurement time down to one breath-hold period, the concept of ‘segmentation’ was introduced. Instead of measuring one

Fourier line per heart beat per cardiac phase, multiple Fourier lines are measured (Fig. 1.62). for example, for a 126*256 matrix and a k-space that is split up into 9 segments, 9 Fourier line per heart beat per cardiac phase, one line for each k-space segment. The measurement time will last one heart beat for the preparation scan, and 14 heart beats are needed to fill up the k-space (14*9=126). Fifteen heart beats are well tolerated for a breath-hold period. The benefit of suspended breathing will result in a crisper representation of the myocardial border, but there is a price to pay. For a gradient-echo sequence with a bandwidth of 195 Hz/pixel, the minimum measurement time for one Fourier line is approximately 9ms. With 9 segments to be measured, this represents a temporal resolution of 81ms. In order to re-establish the temporal resolution, the measurement of one segment, the segment containing the low k-space frequencies, is placed between the measurements of adjacent cardiac phases as illustrated in Fig. 1.63, and while sharing the Fourier lines of adjacent measurements, a true image of another cardiac phase can be reconstructed. This leads to an overall temporal resolution of, e.g., 57 ms.

Fig. 1.62. Prospective ECG-triggered ‘segmented’ acquisition within a breath-hold. For reasons of simplicity, the illustration shows a k-space divided into three segments. One Fourier line is measured per segment per heart beat per cardiac phase. The typical situation would be a k-space divided into 9 segments. For a 126 matrix size, this would require 14 heart beats to fill the 14 Fourier lines within each of the 9 segments (9*14=126).

Fig. 1.63. Prospective ECG-triggered ‘segmented’ acquisition within a breath hold with echo sharing. For reasons of simplicity, the illustration shows a k-space divided into three segments. One additional segment, containing the lower k-space frequencies, is measured in between the k-spaces segments of adjacent cardiac phases (indicated by the arrows). As illustrated, ‘sharing’ the information contained in the adjacent cardiac phases will allow the reconstruction of an additional cardiac phase, leading to an improved temporal resolution.

1.3.3

‘Dark Blood’ Preparation

The inflow of unsaturated blood into the imaging slice or slab has been utilized in MRA to display the vasculature. In cardiac imaging, the hyperintense blood in conjunction with phase changes due to motion and acceleration causes severe flow artifacts in conventional spin-echo imaging. For T1-weighted conventional spin-echo imaging, these artifacts are suppressed with parallel saturation blocks distal and proximal to the stack of slices and with additional dephasing gradients to cause a signal loss for the moving blood. With the introduction of T2-weighted imaging within a breath hold with fast spin-echo, a more sophisticated solution has been presented and dubbed ‘dark blood’ preparation. As illustrated in Fig. 1.64, the sequence starts with a non-selective inversion of all the magnetization, followed immediately by a selective re-inversion for the slice to be imaged. This all takes place with the advent of the ECG signal, where the heart is still in end-diastole. During the waiting period to follow, the re-inverted blood

is washed out of the slice and is replaced by inverted, saturated blood. As the heart is moving into diastole, the TSE acquisition starts, producing ‘dark blood’ images. T2-weighted imaging of the beating heart within a breath hold is a single-slice technique.

1.3.4

Coronary Artery Imaging and the Navigator Technique

Coronary artery imaging started with a single-slice gradient-echo imaging approach, using the same effect as utilized in 2D-ToF-MRA. The sequence was triggered and segmented, and the data acquisition was placed in end-diastole, where the heart moves less and where the flow within the coronaries is supposed to be maximal. The usual degree of segmentation allowed data acquisition within 11 heart beats. A further refinement was implemented using variable flip-angles, that is, low flip-angles for the first Fourier lines to be measured per heart beat followed by increased flip-angles for the sub-sequent k-space segments. This approach compensated for saturation effects during data acquisition. Searching for coronary arteries with a single-slice approach is cumbersome, lengthy, and often not convincing. Questionable areas often remain questionable. A 3D approach is needed for retrospective reconstruction of the coronary vasculature. A 3D approach using similar parameters is too slow to be performed within one breath-hold period. A solution was presented using a so-called navigator technique as illustrated in Fig. 1.65. The ‘navigator’, a sagittal 2D slice or rod is placed through the liver to monitor the position of the liver/lung interface, and a 3D gradient-echo sequence slab is placed where the proximal parts of the coronary arteries are expected. While the 3D data set is executing the measurement of the same Fourier line multiple times, the data for the ‘navigator’ is collected immediately afterwards. If the position of the liver/lung interface indicates ‘close to expiration’, the Fourier line of the 3D set data is accepted, otherwise it is waived. The position of the liver/lung interface can also be used prospectively to correct the position of the 3D slab (prospective acquisition correction, PACE). Doing so will reduce the measurement time since fewer Fourier lines of the 3D acquisition will be waived.

Recent advances in sequence development show that 3D imaging of the coronary arteries can also be performed within a breath-hold, in conjunction with T1-shortening contrast agent.

Fig. 1.64. ‘Dark blood’ preparation scheme. With the detection of the QRS complex within the ECG signal, a nonselective RF inversion pulse is executed, immediately followed by a selective ‘re-inversion’ for the slice to be imaged. During a waiting period, the re-inverted blood is washed out of the slice and is replaced by inverted blood. The TSE image acquisition to follow will produce a ‘dark blood’ image. Shown is a short-axis perspective of the right and left ventricle surrounded by epicardial fat.

Fig. 1.65. The ‘navigator’ technique. A 2D-slice or rod is placed across the liver/lung interface to monitor the respiratory cycle. A 3D-ToF-MRA slab is placed across the coronary arteries.

The position of the liver/lung interface is evaluated, and the 3D slab is moved prospectively and/or is taken as information to reject/accept the Fourier line for the 3D data set.

1.4

Artifacts in Magnetic Resonance Imaging

There are three types of artifacts in MR imaging. The first group is intrinsic to the method or the imaging technique and is almost unavoidable. These artifacts include chemical shifts, the flow and motion artifacts, and the susceptibility gradients within the patient locally destroying the magnetic field homogeneity. The second group of artifacts contains the avoidable artifacts produced by the user him or himself, which are primarily aliasing artifacts. The third group of artifacts includes those based on compromised system design or system malfunction.

1.4.1

Unavoidable Artifacts

1.4.1.1

Chemical Shift

Since the frequency information is used for selective excitation and spatial encoding, the fact that fat and water have a slightly different resonance frequency will lead to a slight shift between the water image and the fat image. A signal void is created at the fat/water interfaces, where only the water is displayed, and the fat signal is assigned to another pixel based on the lower resonance frequency. Usually, a hyperintense rim is seen at the opposite location of the organ in question, where the fat signal is assigned to a pixel already representing a voxel with the associated water content (Fig. 1.66).

1.4.1.2

Flow and Motion

The phase of the macroscopic magnetization is used as spatial information in the direction perpendicular to the direction of frequency encoding. As discussed previously, flow and motion destroy this phase conference. Moving and flowing objects often have the wrong phase position, not corresponding to the phase position of stationary tissue at the same location. In that case, the pixel representing that location will remain black and/or the phase position of the moving object might correspond to a location outside of the body. In that case,the bright