MRA Techniques and Applications in Atlas 2009

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INTRODUCTIONThe intrinsic sensitivity of nuclear magnetic resonance tomotion first was recognized in the 1950s (1,2). From theearly clinical experiences with magnetic resonance (MR)imaging it was appreciated that moving spins exhibited sig-nal variability. Initially this often was considered a nuisance;the resulting signal instability produced confusing and un-wanted image artifacts. Considerable effort was aimed atminimizing these troublesome manifestations (3). Subse-quently, the signal alterations resulting from flowing bloodhave been exploited to provide vascular morphology andquantitative flow data and are the basis for magnetic reso-nance angiography (MRA). MRA with MR provides addi-tional insights into cerebrovascular diseases above that pro-vided by either technique alone.

Conventional MR yields parenchymal and vascularanatomic information with limited physiologic information.The addition of MRA opens the door to anatomic and phys-iologic vascular information often not obtainable solelyfrom other diagnostic techniques. It often is tempting to as-sume that MRA provides the same information as conven-tional catheter angiography (CA) as a result of the similarityin the display. But MRA displays the physiologic conse-quences of flow dynamics in blood vessels, and therefore thetrue anatomic lumen may not always be reflected in theMRA image. The physiologic and morphologic informationthat can be extracted from a particular MRA sequencevaries. It is important to recognize that along with anatomicimages, details of flow direction, velocity, and volume can be extracted using specific MRA techniques. Thus, MR offers the possibility of combined, noninvasive imaging ofparenchymal anatomy, the vascular supply, and a quantita-tive measure of blood flow in a single high-contrast, high-resolution examination.

Although extensive knowledge of MR physics is not nec-essary to interpret MRA images, a basic understanding isnecessary to establish a protocol for an appropriate MRtechnique for a particular clinical situation. Moreover, fun-damental knowledge will allow common diagnostic pitfallsto be avoided. For a detailed discussion of the fundamentalsof flow theory as it relates to MR imaging, see Chapter 4. In this chapter, we provide a clinicians view of the essentialsof MR physics as they relate to MRA and their clinical applications.

BASIC PRINCIPLES OF MR ANDTHE EFFECTS OF FLOW

The basic MR imaging experiment consists in two essential(and relatively independent) components: (i) excitation/satura-tion, from application of a radiofrequency (RF) pulse se-quence, and (ii) signal sampling/localization, forming the MRimage using magnetic field gradients. Blood flow during eitherexcitation or receiving results in two types of correspondingeffects on the MR signal of moving spins: (i) the wash-in/wash-out or flight of spins relative to the timing and placement ofan RF pulse produces so-called time-of-flight (TOF) effects,and (ii) spins moving during the application, and in the direc-tion of, an imaging gradient produce a shift in signal phasethat depends on the type of flow (constant velocity, disturbed,etc.) and gradient in the flow direction, that is, spin-phase phe-nomena (4,5). These flow-induced changes in the MR signalform the substrate for imaging and quantifying blood flow.The moving spin signal changes may be selectively displayedor enhanced as either a magnitude or phase image display.

Time-of-Flight Effects

TOF effects influence the signal intensities of moving blood instandard imaging situations. Flow void (high-velocity signalloss) and flow-related enhancement commonly seen in conven-tional spin echo (SE) and gradient echo (GE) imaging, respec-tively, are described in Chapter 4. TOF effects reflect themacroscopic motion of spins and their state of longitudinalmagnetization. Most commonly, the magnetization of a bolusof blood is modified (with an RF pulse) at one location and de-tected at another. The time elapsed between the RF labelingand sampling of the flowing spins magnetization is referred toas the TOF effect (Fig. 16.1).

Moving spins result in substantial alteration of both theamplitude and phase of MR signal while using conventionalimaging schemes. If the direction of flow is perpendicular tothe imaging slice (or volume), partially saturated spins are re-placed by the inflowing unsaturated (fully relaxed and magne-tized) spins during the RF repetition time (TR) (5) (Fig. 16.1).The displacement of partially saturated spins is directly depen-dent on the flow velocity and slice thickness. If the TR is shortrelative to the longitudinal relaxation time T1 of stationary tissue, the signal of the stationary material is saturated and

CHAPTER 16 MR ANGIOGRAPHY:TECHNIQUES AND CLINICAL APPLICATIONSJOSEPH E. HEISERMAN, THOMAS J. MASARYK, AND NAFI AYGN

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therefore attenuated (5). Signal from the unsaturated movingblood, which is, therefore, not attenuated, entering the excita-tion volume between pulse sequence repetitions consequentlyhas a high signal intensity compared to the surrounding sta-tionary tissue (flow-related enhancement) (Fig. 16.1). Animportant point to remember is that with conventional SE sequences, spins must receive both a 90- and a 180-degreeslice-selective RF pulse to create the echo. Hence, flow signaleventually decreases at higher flow velocities because of the outflow (high-velocity signal loss) that occurs when thespins move outside the imaging slice before application of the180-degree refocusing RF pulse (5).

Spin-Phase Phenomena

Another class of flow effects results from changes in the phaseof the transverse magnetization that occur when the spinsmove along the magnetic field gradients used for position en-coding in MR. The phase change predicted by the Larmorequation for stationary spins is given by

Gt (no flow) [1]

where is the change in phase during time t, is a constantcalled the gyromagnetic ratio, and G is the magnetic field gra-dient amplitude. The dephasing effect of magnetic field gradi-ents on a sample of stationary spins is recognized widely, andtypically it is compensated for by a second gradient (e.g., thefirst inverted lobe of a bipolar read gradient in a GE sequence),resulting in the complete alignment of all spins within an imag-ing volume element at the time the gradient echo is sampled.

In the presence of motion such as flow along the directionof the gradient, the position of a volume element of interestchanges linearly with the velocity v, or, in symbols, x x0 vt.This can be used to modify the phase change relation to ac-count for the motion. It is important to recognize the direc-tional nature of this phenomenon: The motion must be in the

direction defined by the imaging gradient. For constant-veloc-ity flow, the change in phase becomes

G(x0t vt2) (constant-velocity flow) [2]

The amount of time from the start of the gradient until echo for-mation for a time-of-flight sequence is known as the field echotime (FE). For our purposes here, this is approximately equal tothe echo time (TE). For constant-velocity flow, the phase changeduring this time is given by Eq. 2 and has a contribution linearin the time (TE) and another contribution increasing as thesquare of TE. In the presence of acceleration, there is an addi-tional contribution depending on the cube of TE. This patterncontinues (the rate of change of acceleration is known as jerk,and it contributes according to the fourth power of TE). Bloodflowing in vessels (e.g., the cervical carotid artery) has a varietyof velocities (as well as higher-order motion). Furthermore, allof these contributions vary with time according to the cardiaccycle, resulting in fluid motion that varies from point to pointand also at different times. The spectrum of velocities andhigher-order motion within each voxel depicting the vascular lu-men result in a variety of signal phases. In conventional SE im-ages, this results in spatial misregistration in the phase-encodedirection (i.e., ghosting) resulting from bulk phase shifts of spinsflowing along the imaging gradients during the finite time re-quired between phase encoding and signal sampling. In addi-tion, spin dephasing as a result of the range of intravoxel veloc-ities produces a spectrum of phases within the voxel, leading topartial or complete signal loss. This phenomenon is most promi-nent in circumstances in which the flow has a wide variety of ve-locities (e.g., near vessel walls and in areas of disordered flow)and when gradients are applied over a relatively long time pe-riod with relatively large amplitude (i.e., frequency encoding).In summary, the effects of motion in the presence of conven-tional signal sampling gradients occur simultaneously with, butindependent of, TOF effects and are known as spin-phase phenomena. These effects may be observed in conventionalmagnitude-reconstructed images as ghosting and signal loss.

Phase-shift effects are observed for flow in all directions,specifically along the slice-select, phase-encoding, and frequency-encoding gradients. The resulting phase change depends on bothsequence gradient structure (timing, amplitude, etc.) and flowparameters such as velocity, acceleration, and higher-order mo-tion. The terms of Eq. 2 can also be manipulated to modify mo-tion-induced phase changes. Reduction of echo time can result ina geometric reduction in phase change. Said another way, thephase dispersion effects are very sensitive to the duration of TE,and a small reduction in TE can have a dramatic effect. Velocity-induced phase dispersion, present in physiologic situations, canalso be manipulated by employing cardiac gating. As describedearlier, the gradient term G in Eq. 2 can be altered using addi-tional gradients of appropriate magnitude and timing, so-calledrephasing or refocusing gradients (flow compensation or gradi-ent moment nulling) to negate flow-induced phase changes. Ad-ditional gradients can be use to compensate for acceleration andother higher-order motion phase changes. However, implemen-tation of refocusing gradient pulses in a given pulse sequenceprolongs the echo time, which may be counterproductive.

FLOW MEASUREMENTTECHNIQUES

Quantitative flow velocity measurements can be obtained us-ing techniques designed to take advantage of the bulk motionof spins relative to the timing of the pulse sequence (TOF ef-fects) and on the phase shift accrued by spins moving along a

Chapter 16: MR Angiography: Techniques and Clinical Applications 827

FIGURE 16.1. Flow-related enhancement schematic of a cross sectionof a vessel. The relaxed spins entering the image slice result in high sig-nal intensity because of the inflow of fully magnetized blood into theexcitation volume. Signal strength is proportional to the fraction ofthe replacement of saturated spins within the imaging volume. Otherfactors that influence signal intensity include the flip angle of the ra-diofrequency pulse and the repetition time.


magnetic gradient of known amplitude and duration. Consid-erations in the selection of an appropriate technique for a par-ticular clinical application include spatial resolution (i.e., theability of the method to isolate flow through a particular ves-sel), temporal resolution (i.e., the speed of data acquisition,which determines the ability to provide instantaneous mea-surements or mean flow rates), and dynamic range (i.e., therange of measurable flow rates).

Time-of-Flight Methods

Several TOF velocity quantification schemes based on a varietyof pulse sequences have been devised. These generally involvetracking of a presaturated (6) or excited (7,813) bolus; how-ever, because they are not commonly employed for neurovas-cular imaging, these techniques will not be further discussed.

Velocity and Flow Quantification from Measured Phase Shifts

Acquisition Techniques

Spins flowing along a magnetic gradient accrue a phase shiftthat can be employed to measure flow velocity. This isachieved by using sequences with gradient structures of speci-fied amplitude, duration, and polarity, which in the presenceof flow induce predictable phase changes (1423).

Flow Evaluation from Two-Dimensional Phase ReconstructedImages. The implementation of the phase-display (phase re-construction, or zebra stripe) imaging concept allows velocityinformation to be obtained by using an uncompensated gra-dient structure for velocity in either the slice-select or read

direction. Because the flow velocity of blood or cerebrospinalfluid (CSF) is pulsatile, cardiac gating or another form of vari-ability suppression is applied to improve temporal resolution(e.g., systolic vs. diastolic flow). Collecting a two-dimensional(2D) image at a specified time in the cardiac cycle and recon-structing a complex image allows a phase image to be calculatedat each pixel location, velocity being proportional to the phasechanges (15,21,24,25). In the mid 1980s, Bryant et al. had vali-dated the accuracy of one technique of velocity-induced phasemeasurements based on measured phase shifts using modifiedstandard MR gradients and comparing flow measurements ofcarotid arteries to those obtained by Doppler ultrasound (US)(14). More recently, studies have measured flow velocities andvolumes in other cerebral vessels (Table 16.1). Similar experi-ments performed with conventional SE imaging sequences havebeen verified by in vivo electromagnetic flow meters in experi-mental animals, with good correlation (25,26).

The practical limitations of the technique include problemsrelating to cardiac gating. Without gating or if frequent mistrig-gering occurs, pulsatility (ghosting) artifacts will be seen (27). Inaddition, background phase errors related to radiofrequencypenetration, magnetic field inhomogeneities, chemical shift, andoffset (uncentered) echoes occur. Offset echoes present no majordifficulties for SE techniques, but are potentially more trouble-some with more commonly used GE methods. Another, moreprominent phase error results from higher-order motion effectsin turbulent jets, which can be reduced by the use of short echotimes or higher-order-moment nulling (28). Moreover, each se-quence gradient structure must be both optimized and cali-brated to maximize the sensitivity over the desired selected ve-locity range. The dynamic range of the phase-sensitive flowmeasurement techniques is equal to a to (180 to 180 degrees) phase difference. Flow above the selected range(based on the timing, amplitude, and duration of the gradientsin the direction of flow) produces a phase difference greater

828 Part Two: Brain

FLOW VELOCITIES IN NORMAL SUBJECTS

TA B L E 1 6 . 1

Volume flow Mean velocity Peak velocityVessel (ref.) (cm/s) (mL/min) (cm/s)

ArteryRight carotid (32) 352 21 46 4 63 4Right carotid (481) 305 31; 310 34a

Left carotid (32) 342 15 46 4 60 3Left carotid (481) 263 12; 249 11a

Basilar (32) 164 12 40 3 51 4Basilar (481) 201 10; 213 14a

Anterior cerebral (481) (R) 49 4; 49 5a

(L) 41 6; 38 5a

Middle cerebral (12) 150 8.3 69 9Middle cerebral (481) (R) 164 34; 160 35a

(L) 141 12; 147 20a

Posterior cerebral (481) (R) 54 8; 57 5a

(L) 49 7; 46 7a

VeinSuperior sagittal sinus (32) 435 121 13 1Superior sagittal sinus (481) 281 19; 285 19a

Internal cerebral (481) (R) 21 1; 18 1a

(L) 23 1; 20 2a

a Indicates flow measurement acquired with a nongated two-dimensional phase contrast slice.L, left; R, right. Modified from Anderson CM, Edelman RR, Turski PA. In: Clinical magnetic angiography. New York:Raven Press, 1993:135, with permission.


than 180 degrees, resulting in signal loss and aliasing of mea-sured phase changes to falsely low velocity values (28,29). Inthis scenario, the phase difference from the rapid flow willfalsely be assigned to a lower velocity value. Thus, the sametechnique may require several gradient structures applied overdifferent anatomic regions and velocity ranges. Signal loss fromphase dispersion from flow in gradient directions other than thedirection of interest or from higher-order motion (e.g., accelera-tion, jerk) can be suppressed partially by additional flow com-pensation with gradient moment nulling (29). This extends thedynamic range of such techniques, allowing the measurement ofnonuniform flow. The sequence also must be recalibrated whenthe gradient strengths are changed, which occurs when the slicethickness or field of view is altered.

Flow Evaluation from Two-Dimensional Gradient-EncodedImages. Using GE sequences for phase-sensitive flow mea-surement shortens data acquisition time, increases signal con-trast between moving spins and stationary tissues, and thusfacilitates quantitative measurement of rapid flow as well astime-resolved measurements throughout the cardiac cycle(19). Background phase errors are overcome by several meth-ods. The first involves interleaving two scans, in which thegradients are reversed relative to each other, generating phasechanges in the regions where flow occurs (velocity imagingwith gradient-recalled echoes) (30). The subtraction of theraw data of two data sets results in effectively doubling thephase change resulting from velocity while all other spurioussources of phase are subtracted out and leads to a square-root-2 improvement in signal-to-noise ratio (SNR). Alterna-tively, a corrected image can also be obtained by collecting agated cine-mode data set and subtracting the first image fromall other cardiac-gated scans, providing the change in velocityand also eliminating spurious phase changes (24). The formermethod reveals the actual flow velocities (rather than velocitychange) and improves the SNR of the measurement. If thecine data are subtracted from a control scan with full motioncompensation gradients, absolute flow velocities are deter-mined rather than velocity differences (as long as the first im-age is fully compensated; if not, acceleration can cause an er-ror) (31,32). Because the gated 2D phase techniques provideboth magnitude images and phase information, quantitativeflow rates in milliliters/minute can be calculated by multiply-ing the measured cross-sectional area of the vessel on themagnitude image by the velocities derived from the phasedata (31,32).

Flow Rates from One-Dimensional Phase Measurements.One-dimensional (1D) projection techniques also can be usedto extract velocity information from phase data. An advantageof this method is the high-speed acquisition providing excel-lent temporal resolution. One such lD technique, called real-time acquisition and velocity evaluation (RACE), was devel-oped to measure flow perpendicular to the slice (through planeflow) from phase data produced by motion in the slice-selectdirection (20). Data are generated by application of a slice-se-lect gradient followed by a read gradient (perpendicular to thedirection of flow) and with no phase-encoding gradient. TheTE varies from 6 to l0 ms and TR from l0 to 40 ms (20 ms of-ten is used to obtain sufficient sampling). The sequence is con-tinuously repeated throughout the cardiac cycle, capturingflow information with a temporal resolution equivalent to theTR (e.g., on velocity point every 20 ms for the duration of themeasurement, usually 10 to 20 seconds). Hence, this techniquesacrifices 2D spatial resolution for excellent temporal resolu-tion, short measurement time, and no requisite gating. As withall phase methods, the sequence can be calibrated to a velocityrange of interest, and echo times must be minimized to avoidhigher-order motion dephasing.

The measured phase difference determined by the lD meth-ods reflects motion contributions from everything within theprojection. Phase shifts from overlapping arterial, venous,CSF, and moving parenchymal structures are combined, dis-rupting the phase information. Overcoming motion-inducedphase shifts from structures outside the region of interest butwithin the projection is achieved by several strategies, includ-ing spatial presaturation, projection dephasing (applying agradient to suppress stationary tissue), collecting a cylinder ofdata (33), and use of oblique measurements to project the ves-sel of interest free from overlap. After allowances for the phasecontributions of other tissues, the velocity calculated using thismethod represents a projection across the vessel and, there-fore, is an averaged velocity. For this reason, RACE yieldspeak systolic velocities lower than those determined by US, thediscrepancy depending on the velocity profile across the vesselduring peak flow.

Three-Dimensional Time-Resolved Methods. Phase contrastbased approaches are limited by the large amount of data re-quired for full flow quantification. A four-dimensional (threespatial dimensions plus cardiac gating for time resolution) dataset with 1-mm isotropic spatial resolution with reasonable temporal resolution can require 30 to 40 minutes to acquire at1.5 T (34). This disadvantage can be addressed by specially de-signed sequences based on undersampling, for example, vastlyundersampled isotropic voxel radial projection imaging (VIPR)(35). This approach allows submillimeter isotropic voxel datasets to be acquired in approximately 5 minutes. Flow data arenot time resolved; however, mean flow rates in medium-sizedintracranial vessels can then be retrospectively determined withstandard deviations less than 10%.

A second approach relies on time efficiencies made possibleby parallel imaging methods at 3 T. The high temporal andspatial resolution achievable makes possible detailed studies ofdynamic blood flow in the intracranial arteries (36).

Clinical Applications

Clinical implementation of these techniques in normal volun-teers has been reported in the carotid, cerebral, basilar,femoral, and pulmonary arteries, as well as the heart and aorta(14,16,18,19,23,37,34,3841); however, the experience in theevaluation of neurovascular pathology is limited. Increases invelocity have been observed in the carotid artery of normalvolunteers as well as in patients with high-flow cerebral vascu-lar malformations (32). Characteristic velocity profiles alsohave been measured at sites of peripheral vascular stenosis(15,21,22). Studies on patients with moyamoya disease (42)and an anecdotal report of application of these methods to apatient with subclavian steal (43) have also been published. Aswith all quantitative methods, clinical studies proving clinicalrelevance have not been performed.

MR ANGIOGRAPHY TECHNIQUESIn addition to flow quantification, the signal variation fromTOF and spin-phase effects can be enhanced or modified toderive angiographic images.

Time-of-Flight Methods

TOF MRA imaging methods provide vascular contrast basedon tagging of the longitudinal magnetization of spins flowinginto a region of interest, typically through relaxation, inver-sion, or subtraction (4456). Unique to the more commonTOF approaches is the creation of vascular contrast during a



single scan acquisition followed by removal of stationary tis-sue, not by subtraction, but by image postprocessing.


Adiabatic Fast Scanning Technique. The initial use of TOF ef-fects in the production of angiographic images, reported byDixon et al., used a method of adiabatic fast passage that selec-tively tagged common carotid inflow using a separate RF coilplaced low on the neck and the carotid bifurcation image down-stream using a head coil (44). Suppression of stationary tissuesignal was achieved with twister gradient pulses. A twisterpulse is a low-amplitude gradient pulse that causes a 360-degreephase wrap over a relatively large region (i.e., voxel) of soft tis-sue that produces signal loss secondary to phase cancellation.Vessels in this region that are significantly smaller than the voxelsize reduce the phase shift, and signal loss will lessen, thereforeenhancing vascular contrast. Newer variations of this techniqueemploy stationary tissue RF saturation/inversion schemes (45).

Inversion Recovery Technique. Another TOF technique, de-veloped by Nishimura and associates, uses single-slice, thick-slab (i.e., projection) images acquired with a surface coilplaced at the carotid bifurcation (46). Contrary to Dixonsmethod, which used an excitation coil for spin tagging low inthe neck, this method employs a local receive surface coil. Vas-cular contrast is generated by two acquisitions by differentplacement (at and below the carotid bifurcation on alternateacquisitions) of a spatially selective inversion pulse (180 de-grees). The angiographic images are created by subtraction ofthe two data acquisitions. The vascular signal intensity fromTOF effects is maximized by varying the time following the in-version pulse (TI) and cardiac gating. Signal loss from phaseeffects secondary to rapid flow are minimized through use ofextremely short echo times and gating to diastolic signal read-out. Advantages include significant background tissue sup-pression and the use of very short echo time (through asym-metric sampling of signal), permitting the first MRAdemonstration of a significant arterial stenosis.

2DFT Gradient-Echo Time-of-Flight Technique. Inflow of un-saturated spins (flow-related enhancement) provides vascularcontrast of flowing blood in vessels using sequential two-di-mensional Fourier transform (2DFT) GE images oriented per-pendicular to the direction of blood flow (47,48) (Figs. 16.1 and16.2). Minimizing slice thickness results in high vessel contrastwith excellent background suppression and very good sensitiv-ity to slow flow. Partial Fourier acquisition methods and motioncompensation gradients allow for relatively short echo times,minimizing signal loss secondary to motion-induced phase ef-fects. The use of short repetition times allows coverage of largeanatomic regions in a reasonable examination time. Initially, be-cause the data were not displayed in a conventional angio-graphic format, Wehrli and Gullbergs work received little im-mediate attention. The potential of this technique waspopularized when Keller et al. enhanced the acquisition by usingthin 2DFT slices to minimize inflow saturation and displayedthe data in an angiographic format using a maximum intensityprojection (MIP) postprocessing algorithm (49). Sequential con-tiguous 2DFT TOF vascular images with selectively applied pre-saturation pulses have been used to produce both arteriogramsof the carotid bifurcation and the intracranial vessels andvenograms of the dural sinuses (48,50). Acquisition and post-processing of data is rapid and is useful especially when exam-ining slow flow and large regions of interest.

Three-Dimensional Fourier Transform Gradient-Echo Time-of-Flight Technique. An analogous inflow-enhancementbasedTOF angiographic technique uses three-dimensional Fouriertransform (3DFT) sequence acquisitions (5153) (Fig. 16.3).

Flow-related enhancement is maximized by orienting the imag-ing volume perpendicular to inflowing blood, optimizing theTR and flip angle (Fig. 16.4), and using a transmit/receive headcoil to avoid excitation and therefore saturation of flowing spinsoutside the imaging volume. Signal loss from intravoxel phasedispersion is reduced through flow compensation gradients,short echo times, and extremely thin slices (thus small voxels)available with 3DFT imaging. The signal loss from phase can-cellation results from variation of phase across an image voxel.Minimizing voxel size results in a narrower effective range of velocity-induced phase shifts per voxel, thereby helping to

830 Part Two: Brain

FIGURE 16.2. Sequential two-dimensional time-of-flight technique.Axial gradient echo images are obtained in a craniocaudad directionwith a superiorly placed tracking saturation pulse to eliminate venoussignal. The direction of acquisition opposing carotid flow minimizesthe saturation from a previous slice.

FIGURE 16.3. Three-dimensional (3D) time-of-flight technique withramped radiofrequency (RF) pulse. The 3D volume is acquired with aramped RF pulse to avoid progressive saturation of blood as it travelscephalad through the 3D volume. Note that the flip angle of the RF ex-citation () increases across the volume. The superior saturation pulsesuppresses signal from inflowing unsaturated venous blood.


minimize the resultant phase cancellation and associated signalloss. Echo times are minimized by implementing gradientechoes (alleviating the time-consuming 180-degree refocusingpulse in SE imaging) and asymmetric sampling (sampling theecho earlier in the readout period). Flow compensation for ac-celeration or higher-order motion terms technically is possiblewith complex gradient configurations; however, in practice, using compensation gradients for first-order flow (constant velocity) alone in combination with a very short TE generally ismore effective than higher-order flow compensation, whichslightly but detrimentally increases the echo time (5257).

The 3D TOF MRA method excites and images a thickslab or volume of tissue. The 3DFT pulse sequence is similarto 2DFT, with the addition of an extra phase-encoding gradi-ent along the slice-select direction to partition the thick vol-ume into multiple, thin slices or partitions (usually rangingfrom 16 to 128 in number). The application of a 3DFT re-constructs the final image. Using this technique, it is possibleto produce 3D data sets with voxel dimensions of less than0.8 0.8 0.8 mm that are smaller than possible with

2DFT TOF techniques (51,52). In addition, each 3DFT thinpartition benefits from an increased SNR (SNR n1/2 , wheren is the number of 3D slices), improved slice profile, and a re-duction in T2* effects as compared to a 2DFT slice (58). Thecombination of very short echo times and first-order flowcompensation with the small voxel element of 3D imagingmakes this technique relatively resistant to the imaging-de-grading effects of disturbed flow, at the expense of limitedsensitivity to slow flow. Using inflow of unsaturated spins(flow-related enhancement) for vascular contrast in 2DFTand 3DFT TOF methods also allows the selective imaging ofvessels based on the direction of blood flow. This is accom-plished by placing a presaturation 90-degree RF pulse adja-cent to one side of the imaging slice or volume. By saturatinginflowing spins before entering the imaging slice, inflow en-hancement is prevented, and the vessel appears iso- or hy-pointense compared to the stationary background tissue (56)(Figs. 16.2, 16.5, and 16.6). This technique is most useful inallowing visualization of the cervical carotid arteries free ofvenous overlap, selective display of the aorta or inferior vena


A

B

FIGURE 16.4. With gradient echo se-quences the flow-related enhancement rel-ative to the T1 relaxation of the surround-ing soft tissues can be varied throughchoice of repetition time and flip angle.Fast, low-angle shot images (FLASH) for(A) a 100-ms repetition time and (B) a200-m repetition time. The respective flipangles are represented by the values ineach image frame. With increasing repeti-tion time, the flow-related enhancementextends farther along the neck, even asthe echo time increases. Notice that thebackground also increases. (From TopMagn Reson Imaging 1991;3:4, with per-mission.)

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cava, examination of the intracranial venous system, anddocumentation of the direction of flow (49,50,59,60).

Black Blood Angiography. Despite the use of short echotimes, gradient refocusing, and small voxels, the inflow orbright blood MRA techniques still occasionally suffer fromsignal loss in regions of complex flow secondary to superim-posed spin-phase phenomena. This can produce signal lossleading to overestimation of stenoses or signal dropout. Onepossible solution recently has been applied to the carotid

bifurcation, which renders the blood black by presaturat-ing inflowing arterial blood below the region of interest. Inaddition, arterial spins are intentionally dephased as theyflow along the imaging gradients by using relatively longecho times. Projection angiograms can be produced using aminimum-intensity projection method analogous to themaximum-intensity algorithm described for so-calledbright blood methods. In this technique, the presence of com-plex flow actually adds to vessel contrast, possibly allowingmore accurate evaluation of stenoses (61,62). This techniqueis frequently incorporated into MR protocols intended forplaque characterization (63,64,65).

Strategies for Optimizing Performance of Time-of-Flight MRA

The broad goal of TOF MRA is to provide an accurate depic-tion of vascular structures noninvasively, and these methodshave been largely successful in meeting this challenge. The bestresults are achieved when an appropriate method is matchedto each clinical application. Optimization of sequence parame-ters and in some cases use of supplementary techniques canimprove diagnostic accuracy.

Reductions in flow-related enhancement can be though ofas being due to saturation effects, dependent on the T1 recov-ery time of blood, and transverse dephasing, dependent on T2and T2* decay times. Saturation effects are important in thesetting of slow or in-plane flow, and can be minimized by us-ing thinner slices, smaller flip angles, and relatively long TR.Thus 2D TOF excels in the setting of slow flow and is often thetechnique of choice for venous imaging. It also has a role inevaluation of cervical carotid stenosis in detecting slow flowdistal to a high-grade stenosis. Transverse dephasing is mini-mized by small voxels, to minimize intravoxel phase cancella-tion, and especially by reductions in echo time. For TE greaterthan about 1 ms, flow compensation gradients can also reducethis source of signal loss. Because of their intrinsically high res-olution and short echo times, 3D TOF sequences are less lim-ited by this source of signal loss.

Innovative supplementary methods have also been devel-oped to reduce saturation and improve vessel conspicuity rel-ative to background. Second-order chemical shift effects canbe used to suppress background soft tissues (Table 16.2 andFig. 16.7). Spatially varying excitation, for example, with aramped variation in flip angle through the imaging volumecan be used to limit saturation in 3D TOF methods (66,67).Magnetization transfer contrast sequences have been used toenhance the ratio of vessel to background signal intensity(68,66) (Fig. 16.8). Fat suppression can also be incorporatedfor this purpose. Spatially variable RF pulses are designed toincrease linearly along the major axis of flow so a lower flipangle is applied where the flowing spins enter the volume andhigher flip angles are applied distally where the vessel/soft tis-sue contrast would normally be limiting (69,70) (Figs. 16.3and 16.5). Background suppression can be performed simul-taneously with fat saturation and magnetization transferpulses. Compared to the conventional MRA pulse sequences,distal small -vessel visualization is improved because of thebackground suppression and the specialized RF pulses in-tracranially (Fig. 16.8). This is most obvious when intracra-nial studies are reconstructed at matrices of 5122. In the neck,the distal cervical internal carotid arteries are well visualizeddespite severe stenoses at the bifurcations and slow antegradeflow distally.

Because of the small difference in Larmor frequency betweenfat and water resonances, there is a phase shift between the fatand water signal contributions to a given voxel, which varies asa function of TE. Choosing a TE at which fat and water are out of phase can provide another source of fat suppression,

832 Part Two: Brain

FIGURE 16.5. Multiple overlapping three-dimensional (3D) time-of-flight slab technique. This technique incorporates advantages of bothsequential two-dimensional Fourier transform and 3D Fourier trans-form exams. The axially oriented partitions maximize flow-related en-hancement. The smaller 3D slabs allow thin contiguous slices, to avoidflow saturation. Note that the slabs are acquired in a craniocaudad di-rection to avoid saturation effects. With the implementation of aramped radiofrequency pulse, larger 3D volumes are used becauseflow saturation is reduced with this technique. A superior saturationsuppresses venous signal.

FIGURE 16.6. Sequential two-dimensional time-of-flight magneticresonance venography. Coronal acquisition, oblique coronal maxi-mum intensity projection. Excellent slow-flow sensitivity of the two-dimensional time-of-flight technique minimizes this source of signalloss.

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contributing to vessel conspicuity. However, the benefit in back-ground soft tissue suppression is usually more than offset by theincrease in intravoxel dephasing and consequent signal loss inareas of disturbed flow that accompanies the increase in TEabove the minimum value. As a general rule, TE should be min-imized at the expense of other considerations because of therapid increase in intravoxel dephasing that accompanies in-creases in TE.

Great progress has been made in minimizing the effects ofnonuniform or turbulent flow in MRA. However, areas oftight stenosis continue to display flow-induced signal loss at

and immediately following the luminal narrowing. This ef-fect is less pronounced with the TOF 3D and sequential 2Dtechniques than with phase-sensitive methods (55). In addi-tion, the 2DFT slices generally are thicker than the individ-ual partitions of a 3DFT technique, resulting in a largervoxel size in the slice-select direction. Relative to the 3DFTmethod, this produces decreased spatial resolution in theslice-select direction and may lead to signal loss from increased intravoxel phase dispersion (discussed later). Inaddition, the larger slice-select gradients necessary for thesequential 2DFT sequences generally prolong the minimum


SECOND-ORDER CHEMICAL SHIFT EFFECT

TA B L E 1 6 . 2

Chemical shift (ms) for given magnetic field strength

Phase 0.5 T 1.0 T 1.5 T 3.0 T

In 0 0 0 0Out 6.8 3.4 2.3 1.1In 13.6 6.8 4.5 2.3Out 20.4 10.2 6.8 3.4In 27.3 13.6 9.1 4.6Out 34.1 17.0 11.4 5.7

Modified from Keller PJ, Drayer BP, Fram EK. Neuroimaging Clin North Am 1992;2(4):643, with permission.

A B

C DFIGURE 16.7. Second-order chemical-shift artifact seen on gradient echo images results from thetransverse magnetization of both fat and water lying in phase. Because water and fat precess atslightly different rates as dictated by the main magnetic field, there are times during the decay thatfat and water will be in or out of phase. A, B: Axial three-dimensional time-of-flight images at 1.5T with an echo time of 5.0 and 7.0 ms, respectively. Note the orbital and subcutaneous fat high-signal intensity with the in-phase and low-signal intensity with the out-of-phase acquisition withan echo time of 7.0 ms. The lateral projections (C, D) illustrate the image degradation by theoverlapping fat.


echo time relative to comparable 3DFT techniques. The op-timization of vessel signal and reduction of flow artifact hasbeen an area of keen interest, primarily as it relates to short-ened gradient times and reduced spin dephasing (7175).These efforts have been devoted to innovations in both im-age reconstruction and system hardware. The use of conven-tional pulse sequence design and reconstruction techniquesfavors the use of some form of 3D TOF acquisition withfirst-order flow compensation for delineation of the cerebralarterial tree (54,55,71,76,77). Reductions in echo time arepossible in conventional systems with strategies such astruncated asymmetric RF pulses, partial Fourier reconstruc-tion in the read direction, and alternative reconstructionmethods (78).

Improved gradient hardware capability has further en-hanced such sequences, primarily by lowering the minimumTE achievable with a given pulse sequence. The absolute echotime determines signal loss as a result of T2* decay processesas well as to losses incurred because of motion through staticlocal field inhomogeneities. Additional motion-induced de-phasing results from motion through the applied imaging gra-dients, with the severity of the associated signal loss depend-ing on the gradient duration and timing (field echo time) andthe gradient strength (71,72). The FE is specific to each spatialaxis and is determined in each case by the gradient durationalong that axis (79). When conventional 2D and 3D TOFMRA sequences are compared to sequences incorporating thehigher gradient capabilities with ultrashort echo times (butotherwise similar), the greatest single factor affecting the vi-sualization of the simulated carotid stenoses was the fieldecho time (51,57,71,72). Higher-order motion terms are ex-pected within and immediately distal to a high-grade stenosis.This explains the dramatic improvement in the size of thepoststenotic flow void and the visualization of the stenoticsegment itself when standard sequences are replaced by thesequences designed for higher gradient capabilities. As noted

by Evans et al., the field echo time should be the most impor-tant consideration, and the results with longer read fieldechoes tended to parallel the changes seen with longer ab-solute TEs (72). The read field echo is affected by the receiverbandwidth, necessitating the use of longer gradients alongthis direction (as compared to the brief slice-select gradientsapplied in 3D). Consequently, with the exception of those se-quences that do not employ higher-than-zero-order giantmagnetoresistance, the read field echo is the primary limitingfactor for reducing the absolute TE. The addition of first-or-der flow compensation in the read and slice-select directionsproduces a further reduction in size of poststenotic flow voidfor the central and eccentric stenoses despite the longer fieldand absolute echo times.

The third characteristic of note reflects the physics of in-flow-based TOF angiography: Bright signal requires notonly flow, but also unsaturated spins. A disadvantage to theuse of such a wide slab of excitation/saturation is that, de-pending on the speed of flow (i.e., residence time within thevolume) and frequency of RF pulses (i.e., TR), one may losevascular contrast deep in the volume secondary to saturationeffects. It is estimated that total loss of flow-related en-hancement occurs after 10 to 20 excitations (57). Althoughrapid arterial blood flow can retain sufficient flow-relatedenhancement to be successfully imaged over small regions ofinterest (such as the intracranial circulation or cervicalcarotid arteries), slowly moving venous, peripheral arterial,or pathologically slowed arterial flow may become suffi-ciently saturated over even small imaging volumes to pre-vent visualization. Thus, inflow-enhanced 3DFT methodsappear limited to relatively small regions of rapid flow (i.e.,arteries of the head and neck). Care must be taken when us-ing the 3DFT methods to allow for adequate inflow throughchoice of repetition time, flip angle, and volume thickness.An inappropriate selection of parameters may result in afalse appearance of vessel tapering.

834 Part Two: Brain

A, B CFIGURE 16.8. Effect of magnetization transfer. All imaging parameters were held the same. A: Three-dimensional time-of-flight magnetic resonance angiography axial image. B: The same image with appli-cation of magnetization transfer. Notice the superior background suppression of the brain tissue; how-ever, the fat within the orbits and scalp now is accentuated. Note that because of the superiorbackground suppression, the visibility of the intracranial vessels is increased. C: Application of magne-tization transfer and fat suppression. Notice on this image the superior delineation of the vessels result-ing from the superior background suppression from both magnetization transfer and fat suppression.(From Lin W, Tkach JA, Haacke EM, et al. Intracranial MR angiography: application of magnetizationtransfer contrast and fat saturation to short gradient-echo, velocity-compensated sequences. Radiology1993;186:753761, with permission.)

[AU5]


The use of small, stacked 3D volumes (multiple overlap-ping thin slab acquisition [MOTSA], multichunk, multislab)may be the most effective method for reducing the problem ofspin saturation (76,69). In effect, this technique combines the2D advantages of reduced spin saturation and the short TEsand small voxels of the 3D TOF MRAs, at the expense of atime penalty depending on the degree of overlap. Careful pa-rameter selection is needed to minimize discontinuity at slabboundaries, which can produce a characteristic venetianblind artifact. When combined with the spatially variable RFpulses, it is possible to increase the size of the volumes, reducethe degree of overlap between the volumes, and reduce thenumber of volumes necessary to cover the region of interest(80) (Fig. 16.5).

k-Space Manipulation. As noted previously, data collectedat the low spatial frequencies or central phase-encodingsteps are the primary determinants of gross contrast in animage, whereas the data collected at the high spatial fre-quencies primarily determine edge definition and have pro-portionately less impact on contrast. Investigators haveused these principles in MRA acquisitions to improve ves-sel/soft-tissue contrast and reduce flow-related dephasing(8183).

Variation in the TR across K-space has been shown to im-prove vessel/soft tissue contrast and reduce acquisition timewith 2D and 3D TOF acquisitions (84,85). In comparison to2D studies, the 3D TOF MRAs are prone especially to spinsaturation because the spins must flow a larger distance be-fore exiting the imaging slab, typically residing for more thanone TR interval within the imaging volume. Consequently,the moving spins experience multiple excitation pulses, and,similar to the adjacent background stationary tissue, becomesaturated. Varying the TR across k-space is particularly ap-pealing because it not only offers the possibility to limit satu-ration, but also potentially reduces the acquisition time. In areverse centric acquisition (kmax/min to k0), applying a longerTR interval to the lowest 22 (of 32) slice-select phase-encod-ing steps produced the greatest benefit in contrast and flow-related enhancement.

Although the longer TR also permits T1 relaxation of thesoft tissues, it permits a longer, more effective magnetizationtransfer saturation pulse that improves the suppression ofbrain parenchyma by 10% beyond what is possible with theconstant, intermediate-length TR interval that is normally ap-plied. The longer TR interval also allows for an additional fre-quency-selective fat saturation pulse to be applied immediatelybefore the acquisition of the lower 15% of the slice-select spa-tial frequency information (85).

An analogous strategy can be applied in a more time-effi-cient manner if data acquired during different portions of thecardiac cycle are assigned to different segments of k-space(86). A 2D TOF sequence would seem more appropriate inthis setting because of the short acquisition time relative to3D studies, and mild improvements have been describedcompared to the conventional 2D sequence (86). When thelower-spatial-frequency data lines are acquired during sys-tole, the rapid influx of fresh unsaturated spins providesstrong flow-related enhancement in those data lines that pri-marily determine contrast in the image. Edge definition of thesmall vessels is determined primarily by the higher-spatial-frequency data lines that are acquired during the quiescentphase of diastole when flow artifacts should be minimized atthe carotid bifurcation. Intracranially, the flow is more con-stant and should have less of an impact on vessel contrastand definition.

TOF MRA: Artifacts and Limitations. Despite the success ofTOF methods, care is required in the acquisition and inter-

pretation of these studies to avoid errors. Both 2D and 3DTOF are associated with distinctive artifacts. Because thestrong gradient play required to specify the thin slice in 2DTOF limits the minimum achievable TE, signal loss in regionsof disordered flow due to intravoxel dephasing is a charac-teristic of this method. This can obscure focal stenoses butalso can contribute to overestimation of stenosis. In contrast,3D TOF and MOTSA-style acquisitions can achieve shorterecho times and demonstrate less intravoxel dephasing. How-ever, because of the relatively large slab thickness, 3D meth-ods are characterized by signal loss arising from saturation.In the presence of slow flow, signal loss can result in an in-correct diagnosis of occlusion and can reduce flow signalwithin aneurysms. Characteristic signal loss can also be seenin the presence of accelerated flow, such as within tortuousvascular segments. Numerical simulations of flow in vesselsand pathologic structures such as aneurysms have shed con-siderable light on the physical basis of these artifacts(8789). In addition to these sources of error, local magneticfield distortions arising from metal or air or bone and tissueinterfaces can result in local dephasing, sometimes simulatingstenosis (90).

Because the source images for TOF MRA are T1 weighted,other sources of T1 hyperintensity can be mistaken for flow-related signal. Examples include fat, gadolinium contrast en-hancement, and subacute hemorrhage. Because the ratio offlow-related to stationary-tissue signal intensity is greater with2D than 3D TOF, these artifacts are more apparent with thelatter technique. The presence of subacute subarachnoid hem-orrhage adjacent to intracranial arteries can lead to both false-positive and false-negative assessments in the search foraneurysms and vasospasm.

Because of the relatively nonoverlapping nature of the spec-trum of artifacts, 2D and 3D TOF can be considered comple-mentary methods, and, in some applications, such as evalua-tion of the cervical arteries, a combination of the twoapproaches increases accuracy.

Postprocessing and Display. Because background soft tissuesare generally well suppressed, the source data in MRA areeasily postprocessed to provide clinically useful representa-tions of the vascular structures. As a first step, 3D imagingdata are often interpolated or zero filled. Although this doesnot improve spatial resolution, it does minimize partial vol-ume artifacts and can improve diagnostic accuracy (9193).Following this, source data images can be manipulated tomake detection and evaluation of abnormalities easier andmore accurate. Three-dimensional or contiguous two-dimen-sional data constitute a volumetric representation. Such datacan be reformatted into more appropriate planes for evalua-tion. Vascular structures can also be segmented from the sta-tionary tissues and projection images resembling conven-tional angiograms produced. The most popular algorithm,MIP, is produced by casting rays through the data set alongthe desired projection angle and selecting the most intensepixel along the ray, discarding the rest (41,94) (Fig. 16.9).More recently, 3D volume-rendering algorithms have be-come available, and, when used on a workstation, they allowinteractive real-time evaluation (Fig. 16.10). This is particu-larly useful for depicting suspicious areas for more detailedevaluation.

Although these representation methods are often useful andcan improve detection of abnormalities (95100) , postpro-cessing generally involves loss of data and can introduce arti-facts. For example, MIP images often overestimate the degreeof stenosis, partially due to poor detection of signal fromslower flow along the margins of vessels and disordered flowin and around stenoses. Hyperintensity on the source imagesnot related to flowfor example, subacute blood products or



fatcan also appear in postprocessed representations andmimic regions of flow (Figs. 16.11 to 16.13). For these rea-sons, correlation of abnormalities seen on postprocessed rep-resentations with source images improves accuracy. Whenquantitative measurements are made, such as residual lumenor percentage stenosis, highest accuracy results from perform-ing the measurements on source images or pixel-thickness or-thogonal reformats.

Contrast-Enhanced 3DFT Magnetic Resonance Angiography.The signal loss related to saturation in TOF MRA can be ad-dressed by rapidly acquiring T1-weighted images during the bo-lus administration of gadolinium-based intravenous contrastmaterial, known as contrast-enhanced (CE) MRA (101). Athigh intravascular concentrations, most of the intravascular sig-nal results from the T1 shortening of blood (from 1.2 secondsdown to 50 to 100 ms) associated with the contrast. To mini-mize venous signal, accurate bolus timing and very rapid imageacquisition are needed. Several methods for bolus timing havebeen investigated, including timing estimates, test bolus, auto-matic triggering, bolus tagging, and fluoroscopic triggering (Fig.16.14). In the standard methods, images are generally acquiredusing fast spoiled gradient recalled echobased sequences(FSPGR, 3D FFE, FLASH). When these are implemented onMR systems fitted with high-performance gradients, repetitiontimes of 3 ms with minimum echo times of about 1 ms are avail-able. These sequences can be combined with novel k-space ac-quisition and phase-ordering schemes, such as partial Fouriertransform (FT) acquisition, and centric and elliptical-centricphase ordering (102) to achieve total acquisition times rangingfrom 10 seconds to 1 minute, depending on the desired resolu-tion (Fig. 16.15).

Although CE MRA addresses the undesirable saturation ef-fects associated with TOF, it is still subject to signal loss asso-ciated with sources of transverse dephasing. However, this ef-fect may also be reduced when very short TE is possible. Theprimary limitation of CE 3D MRA is the k-space acquisitionspeed and thus the achievable resolution during the time-lim-ited first pass of contrast agent. In addition, the fast spoiledgradient echo sequence is intrinsically limited in SNR becauseof fractional radiofrequency, fractional echo, and increasedbandwidth considerations. Fain et al. analyzed the theoretical

836 Part Two: Brain

A B

A

B

FIGURE 16.9. Maximum intensity projection (MIP). This is the mostcommonly used postprocessing method for magnetic resonance an-giography (MRA). A: For any given projection, rays are cast throughthe data set at the chosen angle, and the pixel along the ray with themaximum intensity is chosen; all other pixels are set to zero intensity.The method works best with data in which vessel signal intensitygreatly exceeds background signal and the signal-to-noise ratio is ade-quate. B: Base projection of intracranial three-dimensional time-of-flight MRA at 3 T. Excellent background suppression results in faith-ful MIP of even small-vessel detail. (Panel A: From Edelman RR,Wentz KU, Mattle H, et al. Projection arteriography and venography:initial clinical results with MR. Radiology 1989;172:351357, withpermission.)

FIGURE 16.10. Postprocessing by volume rendering. This method can be used in real time to study sub-volumes of a magnetic resonance angiography data set in detail. Optimal projection angles often depictsubtle abnormalities to best advantage, as in this subtle posterior cerebral artery narrowing, seen with vol-ume rendering (A) and maximum intensity projection (B).



A B

C

A, B

FIGURE 16.11. Right petrous apex choles-terol cyst. A: Note the high signal intensity onthe T2-weighted spin echo image just poste-rior to the right carotid canal (arrow). B: Ax-ial three-dimensional Fourier transform time-of-flight magnetic resonance angiographydemonstrates the high signal intensity justposterior to the right petrous portion of theinternal carotid artery (arrow). C: The maxi-mum intensity imaging incorporates the highsignal intensity from the cholesterol cyst intothe angiogram, simulating a petrous carotidaneurysm (arrow).

limits of spatial resolution in elliptical-centric CE 3D MRA(103). They demonstrated that maximum attainable resolu-tion relates to the TR, field of view (FOV), trigger time, andbolus profile characteristics. FOV dependence is isotropic in this view order, that is, reducing the FOV in either phase-encoding direction improves resolution in both of the phase-encoding directions. Reduced TR also results in improved

resolution because SNR penalties dictated by decreased voxelsize and short TR at small FOV will be compensated by morefavorable k-space weighting, assuming long scan times with aprolonged contrast enhancement tail.

If images can be obtained very rapidly, on the order of 1 frame per second, then time-resolved MRA becomes possible(104). This allows information regarding the kinetics of

FIGURE 16.12. Catheter digital subtraction angiogramand three-dimensional (3D) Fourier transform time-of-flight (TOF) magnetic resonance angiography (MRA) ofthe carotid bifurcation. A: The catheter angiogram de-monstrates a very severe stenosis (long arrow). The twofacing short arrows demonstrate the normal arterial vesselsegment diameter that would be used for calculation of thepercentage of stenosis using the North American CarotidEndarterectomy Trial criteria. B: Notice that the 3D TOFMRA does not resemble the catheter angiogram becauseof incorporation of high signal intensity from a muralhematoma (hemorrhagic plaque) in the angiogram. This isone of the potential pitfalls of MRA of the carotids. Thistends to be less of a problem with 2D TOF because back-ground soft tissue is more effectively suppressed.


contrast inflow to be obtained, analogous to conventional an-giography. To achieve this, trade-offs are necessary, usually re-sulting in reduced spatial resolution or SNR. Various methodshave been described. In 3D time-resolved imaging of contrastkinetics (TRICKS), the central portion of k-space is acquiredmore frequently than the periphery. Each acquisition of thecenter of k-space is then combined with recent data from theperiphery to produce a time series of images with good spatialand temporal resolution (105). A variant known as TREAT(time-resolved echo shared acquisition technique), employinga less discontinuous mapping of k-space, has also been de-scribed (106,107). In another approach, VIPR, radial sam-pling of k-space is used, with a narrow acquisition window forthe central portion of k-space and a wider window for the pe-ripheral portions (108).

The use of gadolinium-based contrast agents for CE MRAresults in a powerful new tool for vascular imaging. Thesemethods address some of the weaknesses of TOF and also re-sult in short imaging times. This can be beneficial, particularlyin the less cooperative patient, although total imaging timesmay not be significantly reduced when the preparation timefor contrast injection is factored in. However, the use of theseintravenous agents is not without risk. Apart from the poten-tial complications and discomfort of venipuncture, thegadolinium-based contrast agents are themselves associatedwith uncommon associated morbidity, such as anaphylactoidreactions and the recently described nephrogenic systemic

fibrosis (109). For this reason, when evaluating these methodsin comparison to alternatives such as TOF MRA, the addi-tional risk needs to be taken into account.

Phase-Sensitive Methods


Magnitude Subtraction. The use of motion-induced phaseshifts combined with subtraction to produce projective MRAswas initially proposed in 1982 (110). The intravoxel velocitydistribution may be modified by cardiac gating so that signal isalternately sampled in diastole (with a small range of in-travoxel velocities/phase dispersion and high intravascular sig-nal) or in systole (in which case the larger velocity distributionleads to signal loss from increased phase distribution and can-cellation). The signal from stationary tissues is identical in sys-tole and diastole, allowing subsequent magnitude subtractionof such images to yield an angiographic image (111).

Signal variability from flow-induced phase changes can bemodified by several factors. The simplest and possibly most ef-fective way to limit signal loss is to reduce the echo time,which results in an exponential reduction in the higher-ordermotion-induced phase change. Alteration of the gradient con-figuration can be used to reduce signal loss from flow-inducedphase shifts as well, with compensation gradients precisely

838 Part Two: Brain

A B

C

FIGURE 16.13. Acute occlusion of the right internal carotid artery and he-morrhagic infarction of the right lentiform nucleus. A: The T2-weighted spinecho image demonstrates the hemorrhagic infarction within the lentiform nu-cleus. B, C: Three-dimensional Fourier transform time-of-flight magnetic res-onance angiography demonstrates minimal signal in the right carotid siphon,most likely resulting from retrograde flow (open arrow). Note the slightly di-minished signal intensity in the middle cerebral artery branches on the rightwhen compared to the left from the slower flow. In addition, there has beenincorporation of the hemorrhage into the maximum intensity projection al-gorithm (arrow). Notice the patent posterior communicating artery on theright side (arrowhead).


designed to cancel the motion-induced phase change at thetime of signal sampling (112). Images with and without theseflow-compensation gradients can be obtained while maintain-ing other imaging parameters (TR/TE) constant to leave sta-tionary tissue with the same signal, allowing subsequent mag-nitude image subtraction to create an angiographicappearance (17,53,57,74,113,114).

Complex Subtraction (Phase-Contrast). A somewhat dif-ferent use of gradient-modulated phase effects depends on

detection of flow as a discrete phase shift, with subsequentcomplex data subtraction/phase reconstruction to produce thefinal angiographic image. This phase-contrast (PC) tech-nique relies on paired data acquisitions with opposite bipolarflow-encoding gradient pulses, resulting in images with vascu-lar signal approximately proportional to the velocity-inducedphase shifts (115117). As with the other phase-sensitive tech-niques, stationary tissue has identical signal on both acquisi-tions and thus subtracts out. Advantages of these phase-sensitivemethods include their high sensitivity to slowly flowing protons,


A

B

CFIGURE 16.14. Contrast-enhanced magnetic resonance angiography (MRA) acquisition schemes. A: Aseries of fast three-dimensional (3D) measurements is simultaneously acquired with the contrast injection.B: Contrast-enhanced MRA with subtraction. Pre- and postcontrast images are acquired, and stationarytissue is removed by subtraction. C: Fluoroscopically triggered contrast-enhanced MRA. Contrast arrivalis monitored by a two-dimensional (2D) sequence, and 3D sequence is initiated immediately after. (FromLaub G. Principles of contrast-enhanced MR angiography. Basic and clinical applications. Magn ResonImaging Clin North Am 1999;7:783795, with permission.)


such as venous blood, as well as high vascular contrast result-ing from excellent stationary tissue subtraction. Hiehle et al.compared the effect of magnetization transfer (MT) with orwithout ramped excitation in 3D TOF MRA to 3D PC for in-travascular signal versus background (118). That study evalu-ated both large arteries [supraclinoid carotid and middle cere-bral artery (MCA)], as well as small distal branches, and canbe summarized by three findings: (i) a statistically significantimprovement using MT in intracranial TOF MRA, (ii) a smallimprovement using ramped excitation, and (iii) a dramatic im-provement using PC over any combination of MT with orwithout ramped excitation TOF. Therefore, regardless of themethods used to optimize the intravascular signal of 3D TOFintracranial MRA, there is a significant improvement from 3DPC. The gradient-sensitive techniques have the additional ad-vantage of no requisite cardiac gating.

Phase-Sensitive Postprocessing

For sensitivity to flow along all three imaging axes, both themagnitude-subtraction (including rephase/dephase subtrac-tion) and complex phase-subtraction techniques require threeimage sets, each sensitive to flow in each direction (e.g., read-out, slice select, or phase encoding) with at least one additionalreference scan to serve as the mask (119,120). The pairedimages may be integrated into a single acquisition by alternat-ing (i.e., interleaving) phase-sensitive y-steps. This interleavedapproach helps to reduce misregistration artifact secondary tomotion between paired acquisitions. The three subtraction im-ages can be analyzed separately, or they can be combined togenerate the final angiographic image. Unfortunately, the needfor multiple acquisitions prolongs examination time, makingsuch studies susceptible to gross patient motion.

Artifacts/Limitations

Implementation of flow-sensitive gradients requires additionaltime before signal sampling, thus prolonging TE. This is par-ticularly disadvantageous in regions of fast nonuniform mo-tion and in GE sequences that are sensitive to other T2* effectsthat may degrade image quality. The phase-sensitive tech-niques are particularly sensitive to image degradation frompulsatile flow, as well as instrument imperfections such as eddycurrents (117). As with the phase-derived MR velocity mea-surements, signal phase aliasing also is a problem in vessels

with complex or rapid flow. Emerging concepts are addressingsome of the shortcomings of PC MRA. An example is thephase-contrast variant of VIPR (PC VIPR) (121). This ad-dresses the relatively long acquisition times associated withhigh-resolution PC MRA by using a novel radial scheme ofdata collection. Phase-contrast MRA has also been performedwith gadolinium enhancement.

Parallel Imaging

The time required to acquire a conventional MR sequence de-pends linearly on the number of phase-encoding steps requiredto fill k-space. Reducing the number of phase encodes to reducescan time will adversely affect spatial resolution. Parallel imagingaddresses this problem by using information available from mul-tiple-channel receive coils to replace some of the phase-encodingsteps, allowing a reduction of scan time. In essence, the differ-ence in signal received by coil elements at different locationsaround the body part being scanned can be used for localization(122). The cost of this benefit is primarily decrease in SNR dueto the reduction in number of phase encodes, and, ignoring po-tential gains from increased coil efficiency, is proportional to thesquare root of the scan time reduction factor, with factors of twoand four being commonly employed with eight-channel headcoils. In some cases, aliasing artifacts may also be introduced.The processing that uses the information from the differentialcoil sensitivity can be used in the image domain, that is, afterFourier transformation (sensitivity encoding [SENSE], array spa-tial sensitivity encoding technique [ASSET]), or in the frequencydomain before Fourier transform (simultaneous acquisition ofspatial harmonics [SMASH], generalized autocalibrating par-tially parallel acquisition [GRAPPA]) (123). From a clinicalpoint of view, the two approaches yield very similar results.

Because they typically have very SNR, MRA sequences arenatural candidates for application of parallel imaging (124).The benefits can be realized either as improved resolution orreduced scan time. Time-resolved MRA sequences in particu-lar benefit from incorporation of parallel imaging.

3-T MRA

The increasingly widespread availability of clinical 3-T whole-body scanners provides access to a number of benefits for MRimaging. Fundamentally, a move from 1.5 to 3 T results in afactor-of-two increase in SNR; however, a number of other

840 Part Two: Brain

A BFIGURE 16.15. Linear asymmetric (A) and elliptical-centric (B) k-space ordering schemes. (From LaubG. Principles of contrast-enhanced MR angiography. Basic and clinical applications. Magn Reson Imag-ing Clin North Am 1999;7:783795, with permission.)


changes also occur, which may be beneficial or detrimental, de-pending on the details of the imaging sequence of interest.From the point of view of TOF MRA, three changes are ofparticular interest. The increase in SNR can be used to shortenscan time at equal resolution, to increase resolution at fixedscan time, or for a combination of the two. An approximately30% prolongation of T1 for soft tissues (but not for blood) results in improved stationary tissue suppression. Finally, de-pending on the value of TE employed, relative fatwater phasedifferences may provide additional background suppression(Table 16.2). On the negative side, the increase in susceptibil-ity by a factor of four can potentially result in increased arti-facts, particularly near the bone/soft tissue and air interfaces atthe skull base.

Preliminary results comparing intracranial TOF MRA at3.0 and 1.5 T have been promising (125127). The increasedSNR can be used to achieve voxel dimensions less than 1 mm3.Improved suppression of background soft tissues reduces arti-facts on postprocessed projection and rendered images (Fig.16.9). Improved vessel contour rendition and superior small-vessel detection are very noticeable compared to 1.5 T. Al-though potentially a disadvantage at 3 T, preliminary workhas not demonstrated a significant increase in susceptibility ar-tifacts near the skull base at 3 T, possibly due to reduction inintravoxel dephasing associated with smaller voxel size (127).

Imaging options have also been evaluated at 3 T. Magneti-zation transfer contrast improves the (already very good) sta-tionary tissue suppression (128), although the implementationof MT pulses lengthens scan time and so is not implemented inmany 3-T 3D TOF protocols (127). The MT pulse sequence isalso associated with increased RF power deposition and softtissue heating (specified as specific absorption ratio), althoughthe RF power deposition can be controlled by specially tai-lored MT pulse sequences (129). Destructive interference fromthe RF can adversely alter the ramp shape in standard rampedRF excitation for 3D TOF; however, this can be modified tomake this technique also useful at 3 T (130).

The additional SNR available at 3.0 T also makes parallelimaging a natural choice. This can be implemented in such away that near-isotropic submillimeter voxel dimensions arepossible with an acceptable scan time, and further improve-ments in vessel detection and characterization are realized(127,131). The benefits of using a 3.0-T field strength alsolead to improvements in CE MRA, particularly by reducingacquisition times. Increased conspicuity of contrast enhance-ment should be associated with increased intravascular signalcompared to background (132). Feasibility studies and pre-liminary clinical investigations have reported excellent vesselconspicuity with protocols tailored to the 3-T systems(133,134). Parallel imaging at 3 T is of particular interest for

time-resolved CE MRA. Using echo sharing and parallel ac-quisition for intracranial MRA allowed Nael et al. to achieveapproximately 1 mm in plane resolution and 2 seconds intemporal resolution in a preliminary study (135).

Future Prospects

It is important to appreciate that MRA is a relatively youngtechnology; new insights into the mechanisms of flow-inducedsignal changes and more robust techniques with fewer artifacts(not only at the bifurcation but also at the circle of Willis) arenot only on the horizon but also in some instances are currentlyavailable. Advanced-gradient designs and more-sophisticatedacquisition schemes/postprocessing will continue to provide sig-nificant improvement in image quality and diagnostic accuracy.

3-T Intracranial MRA3D TOF MRA can be substantially improved at 3 T

Improved spatial resolutionImproved vessel contrast and background suppression

CE MRA also benefits from very high fieldImproved spatial and temporal resolutionIncreased gadolinium effect at 3 T can result in reduced

contrast volume

CLINICAL APPLICATIONSMRA provides high-quality, reproducible images, as well aspotentially flow measurements that can be applied to manytypes of cerebral vascular anatomy and pathology. Essentialadditional diagnostic information can be obtained in pa-tients undergoing a parenchymal MR, and the availability ofadditional functional information with diffusion, perfusion,and MR spectroscopic studies enables a complete vascularevaluation to be performed. The TOF MRAbased tech-niques are completely noninvasive, rapid, and easily ac-quired as an additional pulse sequence in conjunction with aconventional parenchymal MR examination. Phase-contrastMRA studies have advantages as supplementary studies, no-tably the ability to quantify flow velocities, volumes, and di-rection. Understanding of the relative advantages and disad-vantages of popular MRA methodologies (including TOF,PC, CE, and time-resolved CE) maximizes the diagnostic in-formation obtained for a particular patient. Differing tech-niques often are complementary; combinations of tech-niques provide the necessary data for an accurate diagnosisand treatment planning. The relative strengths and weak-nesses of the most commonly used TOF and PC techniquesare presented in Table 16.3.


COMPARISON OF TIME-OF-FLIGHT, PHASE CONTRAST, AND CONTRAST ENHANCED MRA

TA B L E 1 6 . 3

2D TOF 3D TOF 3D TOF MOTSA 2D PCa 3D PCa CE

Slow flow Fast disturbed flow Resolution Signal/noise Contrast/noise Acquisition time

a Requires a priori selection of velocity to be measured.2D, two-dimensional; 3D, three-dimensional; CE, contrast enhanced; MOTSA, multiple overlapping thin slab acquisition; MRA, magnetic resonanceangiography; PC, phase contrast; TOF, time-of-flight.


Extracranial Circulation

It is beyond the scope of this chapter to give a full discussionof cerebrovascular ischemic disease; however, it is importantto articulate the pathophysiology of cerebral ischemia and in-farction. Several basic mechanisms have been identified thatresult in ischemia and ischemic infarction, including global orlocal perfusion failure (as in systemic hypoperfusion), vascularocclusion from embolic events, or propagation of thrombus.Clinically, characteristic signs and symptoms often l aid in dif-ferentiating the vascular territories involved and aid in direct-ing the MRA exam.

Carotid Territory Atherosclerosis

The interest in carotid imaging relates to the high incidence ofthromboembolic ischemia and stroke in the United States re-sulting from surgically accessible lesions (136143). Athero-sclerotic carotid artery disease results in cerebral ischemia inseveral ways, including decreased blood flow due to stenosesand by thrombotic or atheromatous embolization at sites ofplaque ulceration that may be quite small (142151). Thus,successful thorough evaluation of the carotid bifurcation re-quires both accurate depiction of the degree of luminal nar-rowing and demonstration of arterial wall irregularity.

Clinical Background. In the mid 1980s, prominent cere-brovascular experts questioned the clinical rationale justifyingthe increasing number of carotid endarterectomies performedin the absence of studies supporting beneficial patient outcome(152155). Specifically, the first large, randomized trial pub-lished by the Joint Study of Extracranial Arterial Occlusion re-vealed no difference in total number of strokes and death be-tween the medical and surgical groups if both perioperativemorbidity and mortality were included in the analysis (140).Subsequent studies failed conclusively to support benefit to pa-tients undergoing carotid endarterectomy; however, most ofthe studies were not well designed. Patient selection criteriawere poorly controlled, and wide variation in perioperativecomplication rates was noted (17,154156). Thus, a consider-able amount of controversy was generated because of the nu-merous anecdotal cases supporting the efficacy of carotid en-darterectomy.

Subsequently, efforts to define the role of endarterectomythrough a carefully controlled definitive multicenter clinical trialwere initiated (157). The North American Carotid Endarterec-tomy Trial (NASCET), the Veterans Administration Sympto-matic Stenosis Trial (VASST), and the European CarotidSurgery Trial (ECST) have addressed some of the controversialissues concerning carotid endarterectomy (158161). TheNASCET objective was to determine whether carotid en-darterectomy in combination with the best medical therapy wassuperior to the best medical treatment alone in patients withcarotid stenoses and transient cerebral ischemia or partialstroke (162). The study goal was accomplished by a large, ran-domized, multicenter trial with strict inclusion and exclusioncriteria. Multiple layers of quality control were incorporated inthe study to ensure proper patient selection and to minimize op-erative morbidity. Both randomized and nonrandomized patientdata were collected, and the data underwent careful statisticalanalysis by independent observers.

A segment of NASCET was stopped when it was recog-nized that patients with 70% to 99% diameter reduction ofthe carotid artery who had been treated with carotid en-darterectomy had an absolute reduction in stroke risk of 17%at 2 years compared to medically managed patients (162,163).The benefits were directly related to the severity of the stenosisand were greater for a stenosis of 99% than for 85% or even70% stenosis. One important exclusion criterion was that of a

tandem stenosis, a finding in 2% of this population. In addi-tion, 2.6% of NASCET patients harbored intracranialaneurysms; 16% of these aneurysms were felt to be clinicallyimportant (164).

The role of carotid endarterectomy in symptomatic moder-ate stenoses (50% to 69%) was assessed in the later stage ofNASCET, and a moderate reduction in the risk of stroke wasreported (165). Endarterectomy was not beneficial in patientswith a stenosis of less than 50%. Similar results were found byECST investigators (166), and subsequent analysis of theECST data and meta-analysis of the existing trials demon-strated good agreement (167169)

It is important to recognize that the severity of the stenoseswas carefully calculated from the most severe stenosis demon-strated on a selective catheter angiogram. For the NASCETtrial, the percentage diameter carotid artery narrowing was de-termined in a specific fashion [1 (stenosis diameter/normaldistal lumen diameter) 100], a standard that has beenadopted by the surgical community (170,171) (Fig. 16.12).Therefore, evaluation of both the common carotid bifurcationusing the NASCET criteria in addition to the remaining extra-and intracranial circulation is sound practice (162,163). TheNASCET trial has established a well-defined practice standardfor preoperative evaluation of the carotid bifurcation: angio-graphically demonstrated and specifically measured stenosis inthe absence of significant tandem lesions. Nevertheless, the au-thors of the NASCET report also stressed that the risk of an-giography be considered when evaluating patients for en-darterectomy and suggested the possible use of a noninvasiveexamination to screen potential catheter angiography candi-dates, thus sparing patients with nonsurgical disease this risk(164,172). There have been efforts to establish an entirelynoninvasive approach to the preoperative imaging of patientswith symptomatic carotid arterial disease. One strategy uses amultimodality approach using color Doppler ultrasonographyand TOF MRA (173175). A study reported by Polak et al.was designed to show that US in combination with MRA com-pares favorably to catheter angiography, thus fulfillingNASCET imaging standards for evaluation of bifurcation dis-ease (175). Although the future of MRA for supplantingcatheter angiography is promising for imaging of occlusive dis-ease of the carotid arteries, there remain unanswered ques-tions: Does MRA/US characterize stenoses of 70% to 99%severity, distinguish complete occlusion from severe stenosis,and exclude tandem stenoses as accurately as catheter angiog-raphy (175)? Only a prospective study of symptomatic pa-tients with color Doppler US, MR/MRA, and catheter angiog-raphy will adequately address these issues, and this has notbeen done (163,176). Therefore, the role of MRA in sympto-matic carotid occlusive disease is in flux and continues toevolve, especially as more-robust MRA techniques continue toemerge (163,164,176180).

For asymptomatic carotid artery narrowing, early workconcluded that the risk of carotid endarterectomy in this pa-tient population was greater than that of medical therapy(181). In a the prospective trial Carotid Artery Stenosis WithAsymptomatic Narrowing: Operative Versus Aspirin(CASANOVA), the authors concluded that there is no signifi-cant difference between patients managed medically with dailyacetylsalicylic acid and dipyridamole versus carotid en-darterectomy (182). This study is very complicated, with manyvariables making any solid conclusion difficult (183). The re-sults from the Veterans Administration (VA) AsymptomaticTrial revealed that carotid endarterectomy prevented transientischemic attacks, but no statistically significant benefit wasdemonstrated in preventing stroke or death (183185). Re-cently, data from the Asymptomatic Carotid Surgery Trial(ACAS) (186) showed that stenoses greater than or equal to 60%of luminal diameter benefit from endarterectomy, with a 6%

842 Part Two: Brain


absolute risk reduction over 5 years. The results of this studyhave been debated vigorously, and the American Heart Asso-ciation Multidisciplinary Consensus Group has advised thatthey consider it acceptable to perform a carotid endarterec-tomy for patients who are asymptomatic if the stenosis is 75%or greater (179). They further classified carotid endarterec-tomy for asymptomatic stenosis of greater than 75% as anuncertain indication if the patient is at high risk in terms ofgeneral medical condition. Additional results from the morerecent Asymptomatic Carotid Surgery Trial (187) have not re-solved the controversy surrounding the indications for en-darterectomy in asymptomatic individuals (188). The issueshave been recently reviewed (189).

Advances in endovascular technology and practice havecreated an alternative to the traditional surgical therapy forcarotid stenosis. Although the use of angioplasty and stentingremains an emerging technique in most centers, it showspromise as a less invasive alternative to carotid endarterec-tomy. Recent reviews have summarized the state of practice(190,191)

Technical Background2D and 3DFT time-of-flight magnetic resonance angiography.The anatomic representation by MRA varies widely, depend-ing on the MRA technique employed. The carotid circulationhas been most extensively studied using 2DFT and 3DFT TOFtechniques (52,75). The 2DFT MRA techniques provide thehighest degree of flow contrast and are better for slow flow(i.e., more resistant to saturation effects) and therefore betterat distinguishing severe stenosis from occlusion (Fig. 16.16).Alternatively, 3DFT TOF MRA techniques possess higher spa-tial resolution and shorter echo times, making them more re-sistant to signal loss from disturbed flow at and distal tostenoses (Figs. 16.17 to 16.19). Clinical studies of both tech-niques have shown reasonably good correlation with catheterangiography; however, these studies have varied significantly

in pulse sequence details such as echo time, voxel size, andflow-compensation schemes. Because of progress in the techni-cal aspects of MRA methods, technical parameters continue toevolve, and thus it is difficult to compare studies performed indifferent eras. Moreover, many reasonable ways to character-ize stenosis exist, including techniques adapted from the mea-surement schemes used to quantify conventional angiogramsin the NASCET and ECST trials, as well as other alternativessuch as the common carotid method (192) and, recently, aresurgence in characterization of residual lumen (193).

The anatomic display from 2DFT TOF MRA in severestenoses often is a flow void or apparent vascular interrup-tion. This flow void is also seen with high-grade stenosis in 3DTOF and reflects a lower limit on the degree of stenosis, al-though this limit is sensitively dependent on TE (194197).Based on these considerations, 2DFT TOF MRA has beendemonstrated to be a highly sensitive screening exam; howeverit may not be the most appropriate technique as the definitive(i.e., final) diagnostic study before angiography and subsequentsurgery. Rather, it plays two important roles, first in determin-ing which patients do not harbor significant disease, and henceneed no further diagnostic studies, and second in identifyingvery slow flow in the presence of severe stenosis. Most studieshave compared 2DFT TOF MRA, or US paired with 2DFTTOF MRA, to catheter angiography; there is insufficient datato allow determination of the combined sensitivity, specificity,and accuracy for the evaluation of the carotid bifurcation, andof importance, distal disease (173,174,198204). Review ofthe literature relevant to carotid MRA reveals several generali-ties: When projection images are used for measurements, 2DFTTOF MRA generally overestimates stenoses, with a wide vari-ation in the correlation with catheter angiography. Most of thecomparison studies had similar grading systems for stenoses,but evaluation of the percentage agreement within stenosisgrades revealed some variation. In 13% to 23% of normalcarotids, 2DFT TOF MRA inappropriately demonstrates mildstenoses (related to the normal disturbed flow at the carotid bi-furcation/bulb) (62,205209). However, these recirculation ar-tifacts are generally distinguishable from true mild stenosis


A BFIGURE 16.16. Carotid artery dissection. A: Two-dimensional time-of-flight (TOF) of the left carotid artery demonstrating a patent inter-nal carotid artery. Notice that the size of the internal carotid artery(arrow

MRA Techniques and Applications in Atlas 2009

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