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Magnetic Resonance Angiography Techniques G. Laub1, J. Gaa2, M. Drobnitzky1 1 2 Siemens AG, Medical Engineering Group, Erlangen, Germany Institute for Clinical Radiology, University Hospital of Mannheim, Germany Introduction MR Imaging (MRI) depends on relaxation times T1 and T2 and spin density r. Furthermore, MRI is very sensitive to motion resulting in a variety of flow effects. On one hand, flow effects are responsible for a number of artifacts which can drastically impair the diagnostic value of the images; on the other hand, flow effects can be used advantageously to develop non-invasive techniques for imaging of the vascular anatomy. These techniques are referred to as Magnetic Resonance Angiography (MRA). The techniques most commonly used for MRA can be classified into two major categories: Time-of-flight and phase contrast MR angiography. Both techniques rely on separate physical effects, and will result in images with different information about the vasculature. More recently, the use of contrast agents in combination with ultra-fast T1-weighted imaging sequences has shown significant improvements in the delineation of the vessel lumen. It is important to note that a proper use of MRA techniques and correct interpretation of MR angiographic images requires a knowledge of the underlying physical mechanisms of flow sensitivity in MRI [1, 2]. Inflow Enhancement The basic idea of Time-of-flight MR Angiography (ToF MRA) is demonstrated in Figure 1a. The blood flow is assumed to be perpendicular to the imaging plane (or volume in the case of a 3D study). For repetition times TR shorter than the longitudinal relaxation time T1 of the stationary spins within the slice, the signal will be reduced due to partial saturation effects. Blood flow in the vessel will move spins from outside the slice which have not been subjected to the spatially selective rf pulses into the imaging slice. These unsaturated or fully relaxed spins have full equilibrium magnetization and, therefore, upon entering the slice will produce a much stronger signal than stationary spins assuming that a gradient echo sequence is applied. This effect is referred to as “entry slice phenomenon”, or “inflow enhancement”, or “flow related enhancement”. 68 electromedica 66 (1998) no. 2 a Ds = v · TR TR << T1 saturated spins flow v unsaturated spins vessel S slice 1a 1b Figure 1a Basic idea of ToF MR Angiography. For pulse repetition times TR less than T1 the stationary spins within the slice are partially saturated and, therefore, will produce only little signal. Spins moving into the slice are fully relaxed, resulting in a high signal intensity. Figure 1b Cross-section in the neck. High vessel contrast is achieved due to saturation of stationary spins and inflow of fully relaxed spins. Signal from the veins is suppressed by an additional saturation pulse which is placed on the cranial side of the imaging slice. A typical application of this effect is demonstrated in Figure 1b which shows a thin section acquired in the neck. Due to the relatively large flip angle of 40° in combination with a short pulse repetition rate of 30 msec, the signal from stationary tissue is almost saturated. In general, blood flow signal from the veins can be sufficiently removed by the application of additional saturation pulses which are applied on the cranial side of the slice. A repetition of the same sequence at different slice positions results in a series of images which show the vascular tree as a sum of individual slices. I Imax The amount of inflow enhancement depends on several factors, including tissue specific parameters like T1, sequence specific parameters (flip angle a, and TR), and geometrical parameters like slice thickness and orientation, or blood flow velocity. Postprocessing of MR Angiography Data In principle, with this technique any vessel segment can be imaged by cutting through the vessel perpendicularly with regard to the flow direction. With repetitive increments of the slice position a three-dimensional data set of the complete vascular structure can be measured [3, 4]. For the observer, this form of representation requires experience in order to obtain the correct threedimensional spatial impression. Obviously, postprocessing methods should be used to extract two-dimensional projections of the vasculature from the three-dimensional volume data [5]. With these methods threedimensional perception can be obtained in two ways – by showing a sequence of projective images with different projection angles or by coding of the depth information onto the surface of the displayed objects. Since the surfaces of most vessels are relatively small, the first method – multiple projections with different angles – has proven more useful in practice. The starting point for this method must be a three-dimensional data set in which the structures to be extracted are associated with a characteristic range of signal intensity levels. In this case a projective image can be calculated by penetrating the data volume with a set of parallel projection rays and selecting along each of these rays only the data point that represents the maximum intensity as demonstrated in Figure 2. The inflow enhancement and proper pulse-sequence parameters (flip angle, pulse repetition time, and flow-compensation) assure that the maximum intensity is always associated with a blood vessel, as long as the projection ray intersects at least one [6, 7]. All of the other projection rays will just pick up a background pixel intensity of the three-dimensional data set. As a result, Figure 3 demonstrates a complete projection image at different viewing angle calculated from one single 3D data set. By varying the projection angle multiple projective images can be obtained retrospectively which allow the observer to obtain the correct spatial perception of the Figure 2 Principle of Maximum Intensity Projection (MIP). Along each ray only the pixel with the highest signal intensity I max is projected onto the imaging plane. A A B C B D C E D E Figure 3 Application of MIP post-processing to a 3D data set in the neck. Different views can be calculated retrospectively to create a 3D perception of the carotid arteries. three-dimensional data set. By displaying a number of projections with projection increments of only a few degrees in a rapid fashion, the perception of a continuously rotated object will be generated which allows a correct three-dimensional visualization of such complex structures as a vessel tree. In some instances, particularly for the best assessment of a vessel narrowing, it is recommended to also evaluate the source images in addition to the MIP results. In complex anatomical situations as in the case of the depiction electromedica 66 (1998) no. 2 69 of the renal arteries it is clearly favorable to restrict the MIP algorithm to those parts of the whole data volume which covers the vasculature of interest thereby avoiding projection-related misinterpretations [9]. ToF MR Angiography Techniques ToF MR Angiography can be classified into three major categories as demonstrated in Figure 4. On the left side (Figure 4a), a sequential two-dimensional technique is shown which provides multiple thin sections of the vessels. Alternatively, a three-dimensional technique can be applied. As demonstrated in Figure 4b, the whole volume is excited simultaneously, and will then be subdivided into thin partitions, or slices by using an additional phase encoding scheme in the slice select direction. Unlike in 2D imaging, where the slice resolution is defined by the excitation profile of the radiofrequency pulse, the slice resolution is defined by spatially encoding magnetic field gradients and can be less than 1 mm. Both techniques – 2D and 3D – are currently used in clinical applications. There are specific advantages, and disadvantages related to each of the techniques. While 2D techniques offer a high vessel/background contrast which is pretty much constant over the coverage of the vasculare structure, the vessel/background contrast in 3D techniques is typically lower, and is progressively getting smaller when spins penetrate through the imaging volume. The slab thickness, or vessel coverage in 3D techniques is therefore limited to a distance at which blood signal approaches a steady state signal. Typically, 3D techniques are applicable in combination with fast flow situations, while 2D techniques may be applied for the visualization of slower flow as well. Another aspect in selecting two- or three-dimensional techniques is related to the spatial resolution. In 2D techniques the spatial resolution is defined by the in-plane resolution (FOV divided by the matrix size, or number of lines, respectively), and the slice thickness of the sequence. Typically, in-plane resolution may be isotropic, like 0.8 mm x 0.8 mm, but slice thickness is more than that, like 2-3 mm, resulting in an anisotropic volume data set. Isotropic resolution, i. e. the voxel size has equal length in all directions, can be obtained with 3D techniques, which also offer a better signal-to-noise ratio due to the averaging effect of the phase encoding in the slab direction. Small vessels, as they typically occur in the intracranial vasculature are generally better visualized with a 3D technique, while larger vessel with a predominantly uni-directional flow, e. g. the common carotid artery, may be visualized very well with 2D techniques. Figure 4 Techniques for ToF MR Angiography: sequential 2D (a), volume imaging (b), multi-slab (c) slice # 1 2 3 slab 4 slab 3 single slab (3DFT) slab 2 slab 1 N flow flow a b a small flow c a water large MT pulse Magnetization reduction as a result of magnetization transfer flow macromolecules frequency imaging volume 5a 70 electromedica 66 (1998) no. 2 dipolar interaction between water and macromolecules 5b Figure 5 Techniques to improve vessel contrast in ToF MR Angiography. 5a: Ramped rf pulses are used to partially compensate for the progressive saturation of spins in 3D ToF MRA. 5b: Magnetization transfer effects are used to improve the vessel-background contrast. One of the major limitation in three-dimensional ToF MRA is the loss of vessel contrast as spins are penetrating into the imaging volume. This effect is due to progressive saturation when spins are experiencing the rf excitation pulses in the imaging volume. A technique which is used to address the spin saturation in 3D ToF MRA techniques is referred to as multislab MRA, or MOTSA (see Figure 4c) [8]. Several slabs are used to create sufficient coverage over the vasculature. Each of the slabs is thin enough to avoid significant spin saturation within this slab. All of the slabs are acquired in a sequential mode, i.e. one slab is acquired after the other. The advantage of this technique is that it combines the isotropic resolution capability of the 3D technique with a relatively small amount of spin saturation similar to the 2D techniques. The disdvantage of the multi-slab technique is related to the imperfections of the slice profile. Therefore, all of the slabs need to overlap by about 20-30% to avoid venetian blind effects in areas in between the slabs. A more effective use of the magnetization is possible when using variable flip angles across the slab as shown in Figure 5a. At the entrance plane when spins enter with equilibrium magnetization a relatively small flip angle is used which still provides sufficient signal, with only little effects on the longitudinal magnetization. The flip angle is increased deeper in the imaging volume to compensate for the reduction of the longitudinal magnetization of the spins on their way through the volume, and so forth. In principle, it is possible to shape the flip angle distribution over the entire imaging volume according to the specific flow velocity and vessel coverage. Another improvement is possible by the application of magnetization transfer pulses (MTC) as demonstrated in Figure 5b. The idea is to use off-resonance rf pulses which do not directly affect the mobile protons which are used to create the signal in MRI. Protons with restricted mobility, however, do get saturated, and because of cross correlation or chemical exchange processes the magnetization will be transfered to some biological tissue such as grey, or white matter resulting in a partial saturation. Blood will not be affected by the MTC pulses, and as a result, there will be more contrast between blood and background. Flow-Induced Phase Shift The second class of MR Angiography techniques is based on the changes in the phase of the transverse magnetization. These phase shifts occur when the spins move along the magnetic field gradients which are used for position encoding in 2D and 3D Fourier imaging. For a gradient pulse pair as shown in Figure 6 the flowinduced phase shift FV is given as FV = g • t • A • v t FV = g • A • t • v My flow FV +A -A time stationary Mx Figure 6 Motion along a bipolar gradient pulse pair results in a flow-induced phase shift of the transverse magnetization. Stationary spins will not be affected by the bipolar gradient pulse pair. The amount of phase shifts depends on the area of each gradient pulse, and distance between the pulses. which shows a linear relationship between the phase shift of a moving spin and its velocity. The flow-induced phase shift can be used for flow imaging and flow quantification techniques. Phase Contrast MR Angiography The basic idea of phase-contrast MRA is to acquire two data sets with a different amount of flow sensitivity as shown in Figure 7 [10]. The first data set S1 is acquired with a flow compensated sequence, i.e. without flow sensitivity. The second data set S2 is acquired with a flow sensitive sequence. The amount of flow sensitivity is controlled by the strength of the gradient pulse pair which is incorporated into the sequence. After completion of the measurement, both signals are subtracted to create a complex difference ÆS. The length of the difference ÆS depends on the phase shift FV within each pixel. An image which shows the signal intensity ÆS represents the velocity of spins at each point within the field of view. Both 2D and 3D acquisition techniques can be applied with phase contrast MRA [11-15]. Alternatively, the phase shift FV can be evaluated from the measurement of S1 and S2. As shown above, the phase shift FV is proportional to the spin’s velocity, and therefore the measurement of FV allows the quantitative assessment of flow velocities. The difference signal ÆS has a maximum value for opposite directions of S1 and S 2, i. e. for FV = 180°. This velocity is typically referred to as venc, and depends on the sequence design, i.e. pulse amplitude and distance between the gradient pulse pair which is used for flow encoding. For velocities larger than venc the difference signal is decreased constantly until it gets zero for FV = 360°. Therefore, in phase contrast (PC) MRA it is important to correctly set the venc of the sequence to the maximum flow velocity which is expected during the measurement. electromedica 66 (1998) no. 2 71 My My My FV Mx Mx flow compensated flow sensitive A bipolar gradient pulse pair produces a phase shift which depends on the velocity component along this gradient. For a complete measurement of the flow velocity more measurements are necessary with orthogonal flow sensitization, e. g. image 1: flow compensated image 2: flow sensitization along x (e.g. read gradient) image 3: flow sensitization along y (e. g. phase encode gradient) image 4: flow sensitization along z (e. g. slice select gradient) postprocessing is done in the following way: v read = image 1 – image 2 v phase = image 1 – image 3 v slice = image 1 – image 4 V = = (v2read + v2phase + v2slice) Gslice Gphase Gread Figure 8 Typical sequence diagram for phase contrast MR Angiography. A bipolar flow encoding pulse pair is applied in subsequent measurements in slice, phase, and read direction, respectively. 72 electromedica 66 (1998) no. 2 S2 ÆS FV S1 Mx Figure 7 Basic idea of phase contrast Angiography. The complex subtraction ÆS of two data sets which are acquired with a different amount of flow sensitivity (e. g. flow compensated S1 and flow sensitized S2) produces an image with signal intensities depending on local flow velocities. subtraction Contrast-Enhanced MR Angiography The signal response of a Flash-type sequence as a function of the flip angle a is shown in the following diagram (Fig. 8) for tissues of various longitudinal relaxation times T1 assuming a sequence with a pulse repetition time TR of 10 msec. In contrast-enhanced MR Angiography (CE-MRA), the use of a paramagnetic extracellular contrast agent (such as Gd-DTPA) increases the blood signal [16, 17]. This is due to shortening of the T1 relaxation time of blood after contrast injection. The amount of signal enhancement with T1 shortening is as shown in Figure 9. Depending on the actual concentration of Gd-DTPA the arterial blood T1 can be as short as 50 ... 100 msec, substantially shorter than the T1 of fat. Therefore, blood produces the largest signal, and thus the vessel lumen will be picked up with the maximum intensity program (MIP) to create an MR angiogram [18-20]. Following the intravenous (IV) injection of the contrast agent, it is delivered to the vasculature of interest. Local blood signal is substantially enhanced. This result is optimal when data collection (in particular the center of k-space) occurs right when the contrast agent arrives at the vessels being imaged as shown in Figure 10. Therefore, timing of the scan with respect to the intravenous injection of contrast agent is very important in CE-MRA. In a clinical setting the contrast arrival time is best determined by administering a small testbolus and continuously monitoring for vascular signal enhancement in the region of interest using a very fast gradient-echo technique with fluoroscopic image update. Furthermore, there will be venous enhancement visible when the acquisition window is long enough to catch the venous phase as well. Therefore, the acquisition window is limited to about 10 to 15 seconds in the carotid artery to avoid excessive signal enhancement in the jugular vein [22]. Similar constraints exist for the first-pass bolus width for imaging of the abdominal vasculature [21, 23]. One major technical prerequisites for CE-MRA is the availability of a high-performance gradient system to reach the very short echo and repetition times needed to cover a resonable 3D-volume in the time given by the arterio-venous transit time. At the same time it is mandatory to use CP-phased array coils for complete volume coverage, fitted to the anatomy, to optimize signal-to-noise. The clinical significance of CE-MRA is expected to grow due to a number of reasons: T1-shortening as the source of signal increase is virtually immune against signal loss due to flow-related dephasing especially for in-plane flow as encountered in ToF- and PC-MR angiography. This is clearly advantageous in situations of tortuous vessels and stenotic regions. Due to the fact that Gd-DTPA has no nephrotoxic potential its use in kidney studies is widely accepted. Furthermore it is reported in the literature that the depiction of large aneu- risms is markedly simplified with CE-MRA compared to DSA because filling and subsequent washout of contrast material is not immediate but can be traced over a longer period of time. The main advantage of CE-MRA compared to ToF- and PC-MRA is its intrinsic advantage in acquisition speed, if sufficient strong gradients are available. This results from the fact that the slice orientation can now be chosen independent of the primary vessel direction and must not longer be perpendicular to the vessel. Instead, the image plane can now be chosen to be along any major vessel axis. We consider CE-MRA as an emerging new vascular application of MR. Its specific advantages with respect to conventional DSA is currently been studied at major research centers. Outcome analysis and its potential to reduce costs are currently considered by a growing number of research groups. Figure 9 Relative signal intensity versus flip angle for a spoiled gradientecho sequence (FLASH) for tissues of various T1, and a TR of 10 msec. The flip angle a is selected to produce the largest signal for the short T1 tissues signal intensity TR = 10 ms T1 = 50 ms T1 = 100 ms T1 = 260 ms (fat) T1 = 1200 ms (blood) arterial phase venous phase contrast injection 8 sec 12 sec 16 sec 20 sec 24 sec 28 sec Figure 10 Timing of contrast-enhanced MR Angiography. In this study a series of ultrafast 3D measurements with only 4 sec scan time was acquired right after bolus injection of a single-dose Gd-DTPA. Note rapid filling of the aorta and subsequent filling of the renal vein, as an example. 32 sec electromedica 66 (1998) no. 2 73 74 Figure 11 Dual-phase contrast-enhanced MRA study. The first measurement (left) is done during the arterial phase. Twenty seconds later a second measurement is performed to get an image during the venous phase. Each measurement is done during a breath-hold of 24 seconds. Figure 12 Carotid MR Angiography. Left: 2D-ToF technique. The stripe artifact is due to swallowing right during the acquisition of this particular slice. Right: Contrast-enhanced 3D MRA technique. A 3D volume is acquired with a scan time of only 10 seconds. Due to the short acquisition time the venous enhancement is relatively small. A stenosis is demonstrated in the left internal carotid artery. Figure 13 Thoracic MR Angiography. Two maximum intensity projections are shown from the same 3D data set. A 1.5 mm isotropic resolution is acquired in a single breath-hold of 23 seconds after injection of a double-dose Gd-DTPA. Figure 14 Application of contrastenhanced MR Angiography in the abdominal vasculature. Left: Demonstration of aneurysms in the abdominal aorta. Right: Demonstration of multivessel disease in the aorta and femoral artery. electromedica 66 (1998) no. 2 Literature: General [1] Anderson, C. M., Edelman, R. R., Turski, P.A.: Clinical Magnetic Resonance Angiography: New York, NY; Raven Press, 1993. [2] Potchen, J. E., Haacke, E. M., Siebert J. E., Gottschalk, A.: Magnetic Resonance Angiography: Concepts and Applications. St. Louis, MO; Mosby, 1993. Time-of-flight MRA [3] Masaryk, T. J., Modic, M. T., Ruggieri, P. M., et al: Three-dimensional (volume) gradient-echo imaging of the carotid bifurcation: preliminary clinical experience. Radiology 171: 801-806, 1989. [4] Masaryk, T. J., Modic, M. T., Ross, J. S., et al: Intracranial circulation: preliminary clinical experience with three-dimensional (volume) MR angiography . Radiology 171: 793-799, 1989. [5] Keller, P. J., Drayer, B. P., Fram, E. K., et al: MR angiography with two-dimensional acquisition and three-dimensional display, Radiology 173: 527-532, 1989. [18] Marchal, G., Michiels, J., Bosmans, H., Van Hecke, P.: Contrastenhanced MRA of the brain. J Comput Assist Tomogr 16: 25-29, 1992. [19] Runge, V. M., Kirsch, J. E., Lee, C: Contrast-enhanced MR angiography, J Magn Reson Imaging 3: 233-239, 1993. [20] Lin, W., Haacke, E. M. , Smith, As., Clampitt, M. E.: Gadoliniumenhanced high-resolution MR angiography with adaptive vessel tracking: preliminary results in the intracranial circulation. J Magn Reson Imaging 2: 277-284, 1992. [21] Prince, M. R.: Body MR angiography with gadolinium contrast agents. Magn Reson Imaging Clin N Am 4: 11-24, 1996. [22] Levy, R., Prince, M.R.: Arterial-phase three-dimensional contrastenhanced MR angiography of the carotid arteries. Am J Roentgenol 167: 211-215, 1996. [23] Gaa, J., Laub, G., Edelman, R. R., Georgi, M.: First clinical results of ultrafast, gadolinium-enhanced dual-phase 3D MR-angiography in the abdomen. Fortschr Röntgenstr 169,2: 135-139, 1998. [6] Marchal, G., Bosmans, H., VanFraeyenhoven, L., et al.: Intracranial vascular lesions: optimization and clinical evaluation of threedimensional time-of-flight MR angiography, Radiology 175: 443-448, 1990. [7] Ruggieri, P. M., Laub, G. A., Masaryk, T. M., Modic, M. T.: Intracranial circulation: pulse-sequence considerations in three-dimensional (volume) MR angiography, Radiology 171: 785-791, 1989. [8] Parker, G. L., Yuan, C., Blatter, D. D.: MR angiography by multiple thin-slab 3D acquisition, Magn Reson Med 17: 434-451, 1991. [9] Litt, A. W., Eidelman, E. M., Pinto, R. S., et al: Diagnosis of carotid artery stenosis: comparison of 2DFT time-of-flight MR angiography with contrast angiography in 50 patients, AJNR 12: 149154, 1991. Phase-contrast MRA [10] Axel, L., Morton, D.: MR flow imaging by velocity-compensated/ uncompensated difference images, J Comput Assist Tomogr 11: 31-34, 1987. [11] Dumoulin, C. L., Souza, S. P., Walker, M. F., Wagle, W.: Threedimensional phase contrast angiography, Magn Res Med 9: 139-149, 1989. [12] Pernicone, J. R., Siebert, J. E., Potchen, E. J., et al: Three-dimensional phase contrast MR angiography in the head and neck: preliminary report, AJNR 11: 457-466, 1990. [13] Pelc, N. J., Bernstein, M. A., Shimakawa, A.: Encoding strategies for three-dimensional phase contrast MR imaging of flow, J Magn Reson Imag 1: 405-413, 1991. [14] Hausmann, R., Lewin, J. S., Laub, G.: Phase-contrast MR angiography with reduced acquisition time: new concepts in sequence design, J Magn Reson Imaging 1: 415-422, 1991. [15] Dumoulin, C. L.: Phase contrast magnetic resonance angiography, Neuroimaging Clinics of North America 2: 21-41, 1992. Contrast-enhanced MRA [16] Creasy, J. L., Price, R. R., Prsbrey, T., et al: Gadolinium-enhanced MR angiography, Radiology 175: 280-283, 1990. [17] Chakeres, D. W., Schmalbrock, P., Brogan, M., et al: Normal venous anatomy: demonstrated with gadopentetate dimeglumine in enhanced 3D MR angiography, AJNR 11: 1107-1118, 1990 Author’s address: Dr. Gerhard Laub Siemens AG, Medical Engineering Group Henkestrasse 127 D-91052 Erlangen, Germany electromedica 66 (1998) no. 2 75