Physiologic Evaluation of the Brain with Magnetic Resonance Imaging

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CHAPTER 19 Physiologic Evaluation of the Brain with Magnetic Resonance Imaging

Amish H. Doshi, Pascal Bou-Haidar, Bradley N. Delman

Although the first magnetic resonance imaging (MRI) sequences provided largely anatomic information, over the past 20 years there have been considerable advances in the development of physiologic sequences that better illustrate disease processes. With MRI, radiologists are increasingly able to suggest more specific disease processes with greater accuracy. Perfusion and diffusion imaging allow rapid diagnosis of ischemia and infarction and can identify the so-called ischemic penumbra that may be rescued with aggressive therapy. The integrity of white matter tracts can be established with diffusion tensor imaging (DTI), and the more recent development of tractography allows graphic visualization of these tracts. Cerebrospinal fluid (CSF) flow patterns are demonstrated by performing cardiac-gated scanning to characterize bulk shifts in water. Blood oxygen level–dependent (BOLD) imaging characterizes changes in flow and oxygenation to specific areas of the brain during specific tasks to characterize the site and degree of brain activation. Spectroscopy offers a more specific understanding of metabolites in different areas of the brain and can therefore indicate areas of necrosis, anaerobic metabolism, and regions of accelerated cell membrane turnover or identify specific metabolites characteristic of specific disease processes.

Diffusion-Weighted Imaging

Diffusion-weighted imaging (DWI) has become an important tool in routine evaluation of the brain on MRI. On modern scanners a diffusion sequence may take 30 seconds or less, so many sites have added this technique to a standard brain-imaging protocol. Although DWI had initially been used predominantly to identify areas of acute ischemia, the appearances of other specific disease entities have been well characterized with this imaging method.

DWI characterizes differences in the brownian motion of water molecules, depending on their local environment. In bulk, water molecules can move freely in any direction. Biologic systems typically impede the free motion of water in one or more directions. Thus, cellular structure, permeability barriers, and various macromolecules within the brain parenchyma may all restrict the free diffusion of water molecules. Normally, restriction tends to be greater within the intracellular space than within the extracellular space ¹; thus, extracellular water molecules can diffuse more freely than intracellular molecules.

Physics

The imaging technique most often used for DWI is an echo planar imaging (EPI) sequence. In EPI, two separate and equal magnetic gradients are applied to opposite sides of the radiofrequency (RF) pulse during image acquisition. Use of these bipolar pulsed gradients allows detection of diffusional motion by changes in the magnitude of the moving spins from phase dispersion. Water molecules that do not travel significantly between excitation pulse and read pulse will not dephase significantly and will therefore retain much of their initial signal. Thus, in pathologic states, the increased signal on DWI reflects an abnormal decrease in water diffusivity, typically caused by either loss of normal extracellular space (as seen in cytotoxic edema) or the presence of a highly viscous, proteinaceous, or cellular environment.^1,² It is worth noting that loss of signal is not due to travel of water from one imaging voxel to another, but rather motion within the voxels themselves. Indeed, the average motion of a water molecule at 40°C is 2.5 × 10⁻³ mm²/sec, which translates into motion of 22 µm in 100 msec; this is considerably smaller than the typical 1- to 2-mm voxel dimension.³

The magnitude of diffusion sensitization of DWI is determined by the “b-value,” which in turn is related to the duration, strength, and time interval between the magnetic gradients. Therefore, higher b-values confer more sensitive DWI and thus yield improved contrast and ability to identify areas of water restriction; however, higher b-values also result in loss of signal, as well as noisier images, which may then have reduced utility. The b-value that is typically used in clinical assessment is approximately 800 to 1000 sec/mm²,⁴ but with certain applications, such as vertebral imaging or characterization of intracranial tumors, b-values may range from 500 to 2000 sec/mm² or greater.

The rate of diffusional motion is characterized by the apparent diffusion coefficient (ADC). This quantitative estimate of diffusivity can be achieved by acquiring two sets of images with different b-values, which can eliminate the effects of spin density and T1 and T2 relaxation. In clinical practice, one of the b-values would measure approximately 1000 sec/mm², as indicated earlier. The other b-value is typically 0 sec/mm², which reflects an image that does not have any motion-probing gradient applied (yielding so-called B0 images). ADC maps quantify differences between B0 and B1000 images and therefore nullify “T2 shine-through,” a condition in which high B1000 signal is primarily due to high T2 signal rather than water restriction (Fig. 19-1). Thus, true restricted diffusion, as seen with acute ischemia, will demonstrate high signal on DWI and corresponding low signal on ADC maps.¹

FIGURE 19-1 Breast carcinoma metastasis producing a T2 shine-through artifact. A, Axial diffusion trace image demonstrating a rounded focus of increased diffusion signal in the right posterior frontal lobe (asterisk). B, Corresponding T2 image (B0 image) showing increased signal in the same distribution (asterisk), thus suggesting that the diffusion signal may not reflect true restriction. C, Apparent diffusion coefficient map confirming that this area is not dark, so the lesion does not truly restrict diffusion.

Clinical Uses and Applications

In clinical use, diffusion images are acquired in at least three directions, with the resulting images mathematically averaged to generate the trace image in which regions of normal white matter and gray matter appear fairly uniform. The normal brain shows slight variation in diffusion signal based on cellular structure. Gray matter has marginally higher signal on DWI, probably reflecting differing T2 properties rather than true differences in the ADC. Areas that are essentially water (most notably CSF) are dark on DWI sequences because of the free motion of water molecules in bulk water.⁵

The most completely characterized application of DWI is for the diagnosis and management of patients with acute cerebral ischemia. When perfusion does not meet the metabolic demand of territorial parenchyma, energy-dependent sodium-potassium adenosine triphosphatase ion pumps may fail. The resulting ion flux leads to the accumulation of water within the intracellular compartment (cytotoxic edema). Neuronal swelling crowds the extracellular space, with resulting restricted motion of extracellular water molecules that is manifest as high DWI signal; these changes may be seen within minutes of tissue infarction. Normal brain ADC values range from approximately 740 to 840 × 10⁻⁶ mm²/sec, with considerable overlap between white matter and gray matter values.⁶ Parenchyma with ADC values of less than 500 to 550 × 10⁻⁶ mm²/sec is almost certainly in the process of irreversible infarction. Intermediate ADC values (550 to 700 × 10⁻⁶ mm²/sec) may reflect parenchyma that is ischemic yet still viable; it is this tissue that may be potentially salvageable with appropriate intervention. Although DWI is frequently identified as the most sensitive sequence for detecting infarction, other complementary sequences (such as the perfusion sequences described later) can help define an apparent impending infarction and define tissue at risk for infarction, the so-called ischemic penumbra.^7,⁸

As infarcts age, the ADC and DWI patterns mature at differing rates. Although both become abnormal within minutes, the ADC increases from low signal to isointensity more rapidly, in most instances within 5 to 14 days. The ADC then continues to rise over the ensuing weeks to remain positive for the life of the patient, corresponding with encephalomalacia. DWI normalizes more slowly. It may reach isointensity with surrounding brain as long as 4 to 5 weeks after the infarction, and it will take even longer for signal to decline further in the setting of encephalomalacia (Fig. 19-2).

FIGURE 19-2 Acute left lenticular infarct. A, Axial non–contrast-enhanced computed tomography (CT) showing a focal chronic lacunar infarct (arrowhead), but no area suspicious for acute infarction. B, Axial T1-weighted image obtained 3 hours after CT showing no definite infarction. C, T2-weighted image showing an area of vague lenticular edema (arrow) distinct from the chronic lacunar infarct (arrowhead). D, A fluid-attenuated inversion recovery (FLAIR) T2-weighted image also shows the lenticular edema (arrow). The chronic lacunar infarct (arrowhead) is less conspicuous on FLAIR than on T2 imaging, probably related to the canceling effect of volume loss and gliosis on this sequence. E, Diffusion-weighted imaging showing increased signal corresponding with the infarct (arrow) in C and D, whereas the chronic focus (arrowhead) has no restriction. F, The apparent diffusion coefficient map is dark in the area of edema, thus confirming true restriction of diffusion (arrow). The lacunar infarction is mostly bright, which suggests facilitated diffusion, as is seen with volume loss (arrowhead).

Certain intracranial neoplasms demonstrate restricted diffusion. Many of these neoplasms are tumors that are high in cellularity, with resulting high signal on DWI and low signal on the ADC map (i.e., true diffusion restriction). Examples include lymphoma, medulloblastoma, and portions of high-grade gliomas. Some tumors demonstrate restriction without high cellularity. Hence, although both epidermoid cysts and arachnoid cysts closely follow fluid intensity on T1, T2, and fluid-attenuated inversion recovery (FLAIR) sequences, the high DWI signal and heterogeneous low signal on the ADC map conferred by desquamated debris allow confident diagnosis of an epidermoid. Arachnoid cysts, which instead contain simple fluid, demonstrate facilitated diffusion and therefore have the opposite diffusion pattern (low DWI signal and high ADC signal).⁹

Brain abscesses are typically peripherally enhancing lesions that are usually surrounded by significant vasogenic edema. Occasionally, these lesions can be difficult to differentiate from other peripherally enhancing masses such as necrotic tumors. The central cavity of an abscess demonstrates high signal on DWI and low signal on ADC maps. The restricted diffusion is most likely attributed to the high viscosity of proteinaceous fluid and the hypercellularity of inflammatory cells.¹⁰^–¹²

Seizure activity can induce changes in water diffusivity because of cellular swelling and fluctuation in extracellular fluid. Cortical restricted diffusion has been seen with prolonged seizure activity, presumably caused by an imbalance between oxygen delivery and consumption.¹³ Focal parenchymal changes in the postictal state can be seen as hyperintense signal changes on T2-weighted sequences with variability in DWI signal. In the setting of ischemia related to seizure activity, DWI may initially demonstrate restricted diffusion as a result of cytotoxic edema, but the restricted diffusion can later progress to decreased diffusion signal, which may represent gliosis of the involved tissue.¹⁴

DWI also aids in the diagnosis and further characterization of other disease processes. Restricted diffusion can be seen in zones of active demyelination in such disease processes as multiple sclerosis and progressive multifocal leukoencephalopathy (Fig. 19-3), in areas of active inflammation typically seen in the medial temporal lobes in herpes simplex encephalitis, in the anterior deep gray nuclei of the basal ganglia and along a discontinuous cortical ribbon in Creutzfeldt-Jacob disease, in some phases of hemorrhage, and in some encephalopathies and leukodystrophies.^1,⁹

FIGURE 19-3 Forty-two-year-old man infected with human immunodeficiency virus and progressive multifocal leukoencephalopathy. A, Axial fluid-attenuated inversion recovery T2-weighted image showing a nonspecific area of high signal, possibly edema, near the vertex (arrow). B, Diffusion-weighted imaging showing a small area of corresponding signal abnormality, with restricted diffusion noted preferentially at the periphery (arrow). C, A slight increase in signal on the apparent diffusion coefficient map suggests T2 shine-through (arrow). Ordinarily, progressive multifocal leukoencephalopathy does not enhance, but with the polydrug regimen and immune reconstitution, one may see more distinctive restriction of diffusion and enhancement, both of which are absent in more indolent cases.

As indicated earlier, facilitated diffusion can be seen in patients with encephalomalacia and arachnoid cysts, with resulting low signal on DWI and high signal on ADC maps. Other entities with a similar pattern include cystic lesions, vasogenic edema, transependymal spread of CSF in hydrocephalus, tumor necrosis, and radiation necrosis.

Pitfalls and Limitations

EPI techniques allow ultrafast acquisition times, thereby almost eliminating motion artifact. However, the long echo train lengths needed to obtain the data render these sequences sensitive to both chemical shifts and magnetic susceptibility. Lipid suppression can be used to resolve some effects of chemical shift. Eddy currents result from the use of rapidly alternating gradients; these currents may result in significant distortion or misregistration between directional acquisition, which translates into blurring and loss of soft tissue contrast in the resulting trace image. Traditional DWI is prone to areas of susceptibility, most notably about the skull base, paranasal sinuses, and petrous bone, where air-bone-tissue interfaces are present.^5,^15,¹⁶ In addition, the presence of paramagnetic and ferromagnetic material such as blood products or metal from surgery, trauma, or dental hardware can produce significant artifact, seen as ghosting, image distortion, and susceptibility artifact (Fig. 19-4). Newer techniques (periodically rotated overlapping parrallel lines with enhanced reconstruction [PROPELLER], BLADE, and others) use a rotating acquisition frame to minimize the effect of such artifacts and increase the sensitivity for abnormalities in previously challenging areas; however, these sequences may be associated with a significant increase in scanning time.

FIGURE 19-4 Axial trace diffusion image showing susceptibility around an arteriovenous malformation. This lesion had been embolized previously. There is significant signal loss within the embolized bed (asterisks) and a halo of high signal distortion around the margin of the embolized nidus (arrows). Note that posteromedially, where a small amount of patent nidus remains, one sees low signal flow voids but considerably less susceptibility and distortion.

Diffusion Tensor Imaging and Tractography

Physics

DTI, a refinement of DWI, characterizes the dominant vector of water motion within voxels. In contrast to DWI, where typically only 3 directions (motion-probing gradients) are obtained, DTI requires at least 6 directions to resolve the mathematical ambiguities related to oblique fiber trajectories. Some centers increase signal either by acquiring multiple acquisitions in these 6 directions or by acquiring 25 or more directions; however, with both techniques, the additional scan time increases the risk for patient motion during the sequence and subsequent misregistration of the data.

As mentioned in the discussion on DWI, the water motion illustrated on diffusion sequences is predominantly extracellular, and the vector of motion tends to parallel the white matter tracts. In DTI, the acquisition of at least six directions allows mathematical construction of a tensor ellipsoid whose major axis points in the direction of the dominant fiber tract within a voxel.¹⁷ In voxels in which white matter tracts are homogeneous and nearly colinear, this ellipsoid is thin and elongated in a so-called prolate or cigar-shaped configuration, with the dominant axis (major eigenvector) being significantly greater than the two perpendicular axes (minor eigenvectors). Such ellipsoids are anisotropic; that is, they represent voxels whose fibers have a strong directional bias. When white matter tracts are not as colinear (i.e., when there are many crossing or divergent fibers) or in areas of gray matter or CSF, the tensor ellipsoid is nearly spherical. These ellipsoids are further characterized as isotropic; that is, they lack significant directional bias.

Once tensors have been calculated, they may be presented for interpretation in numerous ways. A mean diffusivity map portrays average molecular motion that is independent of tissue directionality, so voxels with high motion such as water will be bright and voxels with low motion such as gray matter will be dark. A fractional anisotropy (FA) map represents the colinearity and integrity of fibers, and therefore large tracts with parallel white matter bundles, as seen in the callosum fibers and descending corticospinal tract, appear white (FA approaches 1). CSF, gray matter, and crossing white matter tracts all lack parallel axonal configuration and thus appear dark or black (FA approaches 0). Color-encoded directional maps overlay fiber directional data on an FA image, with typical color assignments of red representing transverse fibers, green representing anteroposterior fibers, and blue representing superoinferior fibers; brighter or more saturated fibers indicate greater colinearity and hence greater anisotropy (Fig. 19-5).

FIGURE 19-5 Normal diffusion tensor maps. A, Color-encoded diffusion tensor map showing primary tract orientations. By convention, red indicates right-left fibers, green indicates anterior-posterior fibers, and blue indicates superior-inferior fibers. Oblique directions are represented by intermediate colors, so for example, as splenial fibers leave the midline and angle toward the occipital lobe, they change from red to orange to yellow to lime-green to green. B, Fractional anisotropy map indicating the degree of anisotropy within given voxels. The voxels containing nearly colinear fibers (e.g., genu, splenium, posterior limb of the internal capsule) appear bright, with fractional anisotropy approaching 1. The gray matter has proportionately fewer fibers that are more arbitrarily oriented; they appear dark and have fractional anisotropy approaching 0. Note that when fiber tracts cross, for example, lateral to the posterior margins of the ventricles, signal drops completely because conventional acquisition and processing techniques are unable to resolve discordant orientation data.

Tractography is a mathematical method of tracking theoretical fiber trajectories by using vector data to incrementally advance to adjacent voxels. Tracking is typically initiated by placing a source “seed” over an area of interest. A second “target” seed can be placed to determine a destination focus to which the trajectory must run, or the second seed may be omitted to allow tracking software to course spontaneously to determine all potential tracts leading from the source seed.

Clinical Uses and Applications

DTI can be a very useful tool for demonstrating the integrity or compromise of tracts running between different areas of the brain. Four classic patterns of abnormal fiber orientation have been described: (1) deviated but otherwise preserved fiber tracts, (2) edematous tracts with diminished FA but preserved fiber tract orientation, (3) infiltrated tracts with diminished FA and abnormal fiber tract orientation, and (4) frank destruction of fiber tracts with negligible anisotropy.¹⁸ Figures 19-6 and 19-7 illustrate varieties of tract disruption.

FIGURE 19-6 Infiltrating parietal glioma. A, Axial T1-weighted image showing a slightly heterogeneous low-signal mass within the left parietal white matter (arrows), with the mass causing effacement of the overlying sulci. The intraparietal sulcus is compressed and displaced anteriorly (black arrowhead). The normal right intraparietal sulcus is shown for reference (white arrowhead). B, Fluid-attenuated inversion recovery T2-weignted image showing high signal corresponding with low intensity on T1 (arrows), and there is little enhancement after gadolinium (C) (arrows) . D, Fractional anisotropy map showing loss of anisotropy corresponding to tumor (arrows). E, Tensor color map demonstrating anterior displacement and more strongly inferior (more blue) orientation of fibers within postcentral gyral white matter (black arrowhead) than in the normal position on the contralateral side (white arrowhead). Note that the normal white matter anatomy seen posteriorly on the right (hatched arrows) is obliterated on the left. F, Diffusion tractography showing displacement of fibers around the tumor (arrows).

FIGURE 19-7 Diffusion tensor characterization of white matter integrity around a glioma centered near the sylvian point. A, The lesion itself has low anisotropy (asterisk) because of its high cellular nature and lack of organized white matter tracts. The surrounding white matter is splayed and compressed (arrows). As the compressed arcuate fasciculus extends posteriorly around the mass, the color orientation changes from green (anteroposterior) to yellow (anteroposterior oblique) to orange (transverse oblique) to red (transverse); as these fibers project anteriorly again, the colors repeat in reverse. B, Tractography demonstrates that the corticospinal tracts are deviated medially (arrows) but appear grossly intact.

Such assessment of tracts can aid in planning treatment and determining prognosis, thereby allowing surgeons and radiation oncologists to protect tracts that are preserved while more aggressively addressing tracts that have already been destroyed. This has tremendous surgical importance because with these data a surgeon can maximize resection while minimizing risk to adjacent structures. A study by Kikuta and colleagues involving patients who had undergone resection for arteriovenous malformations (AVMs) found that incomplete tractography of the optic radiation was associated with visual field loss postoperatively.¹⁹ Such information not only assists in surgical planning but also helps prepare patients for potential outcomes of surgery. In addition, for patients undergoing surgery to treat brain tumors near the motor pathways, tractography may be used to identify initial sites for electrocortical stimulation, thus facilitating faster localization of eloquent cortex during surgery.

Applications of DTI extend beyond tumor evaluation. Preliminary studies suggest that DTI may be sufficiently sensitive to detect early changes in vulnerable regions in individuals with certain cognitive disorders, even at the presymptomatic or preclinical stages.²⁰ Additional uses for tractography include assessment of patients with multiple sclerosis, and early studies in stroke patients suggest that DTI can aid in prognosis and may detect corticocortical rewiring.

DTI and tractography have recently been applied to the spinal cord, an area previously too degraded by artifact to image successfully. Some authors now suggest that FA data from DTI may be useful in distinguishing surrounding cord edema from tumor.²¹ In spinal cord AVMs, FA values in surrounding cord may improve after embolization and correlate with enhanced patient outcome.

Pitfalls and Limitations

A major limitation of conventional DTI is its poor distinction of crossing fiber tracts. When a voxel contains a homogeneous population of similarly oriented fibers, calculation of the dominant vector or water motion is straightforward. However, when a voxel contains crossing fibers, a simple tensor calculation is unable to reflect the more complex fiber configuration. In this instance, the major eigenvector may be calculated as the average of all voxel fibers, and as a result the eigenvector may not represent any of the dominant tracts.²² Newer techniques such as high–angular resolution diffusion imaging (HARDI) and Q-ball imaging may use 100 or more directions along with complex mathematical techniques to better characterize intravoxel tract ambiguity.²³

The DTI sequence is also prone to artifacts that result in local field distortion and sometimes signal dropout. Thus, the diffusion signal may be compromised in regions adjacent to postoperative air, blood products, and surgical clips or embolic material. As in DWI, eddy currents can also distort images in DTI and result in image misregistration and computational errors; in these instances, correction software can compensate for at least some of the distortion to yield more accurate maps.

Tractography depends on user input to define tracts of interest. Incorrect or suboptimal region-of-interest placement will result in misrepresented tracts. Mathematical assumptions will also influence tractography generation, with tract validity being dependent on appropriate definition of the maximal angulation within a voxel, the minimum FA value to tolerate while generating the tract, and the total path length. Because signal in the cord is more variable than signal in the brain and because the cord moves with CSF pulsation, spinal tractography proves considerably more challenging. In general, tractography in the cervical cord is easier to generate than in the thoracic cord, at least in part because of the greater pulsation effects in the mid and lower cord.²¹

Magnetic Resonance Angiography

Evaluation of the intracranial vascular system is important in the diagnosis and planning of treatment of many vascular-related disease processes such as aneurysms, AVMs, infarction, and sinus venous thrombosis. Although the “gold standard” for evaluation of many vascular entities is currently digital subtraction angiography (DSA), magnetic resonance angiography (MRA) offers an alternative for the assessment of intracranial vessels that is both noninvasive and does not require the use of ionizing radiation. The three main varieties of MRA used clinically are time-of-flight (TOF), phase-contrast (PC), and contrast-enhanced (CE) techniques. Each MRA is typically postprocessed by a technologist before interpretation to yield maximum intensity projections that can be rotated in space for a three-dimensional (3D) effect (Fig. 19-8).

FIGURE 19-8 Collapsed maximum intensity projection of a three-dimensional time-of-flight magnetic resonance angiogram. Although the technique is optimized to evaluate for visualization of arteries, veins with high flow are also seen, including the sigmoid/jugular systems and sphenoparietal drainage. A1, first segment of the anterior cerebral artery; A2, second segment of the anterior cerebral artery; ACom, anterior communicating artery; AICA, anterior inferior cerebellar artery; ATA, anterior temporal artery; BA, basilar artery; CeICA, cervical internal carotid artery; cICA, cavernous internal carotid artery; eVA, extracranial vertebral artery; M1, first segment of the middle cerebral artery; OA, ophthalmic artery; P1, first segment of the posterior cerebral artery; P2, second segment of the posterior cerebral artery; pICA, petrous internal carotid artery; PICA, posterior inferior cerebellar artery; purple arrowheads, sylvian middle cerebral artery branches; SCA, superior cerebellar artery; sICA, supraclinoid internal carotid artery; VA, intracranial vertebral artery.

Physics

TOF MRA is performed by applying repetitive pulses to stationary tissues in a discrete volume of the brain, which results in saturation or suppression of signal in these tissues. Blood flowing into the volume has not been saturated and is therefore fully magnetized; this in-flowing blood therefore provides the only significant signal within the imaging slab. The intrinsic contrast that is derived from flowing blood eliminates the need for injection of contrast material to visualize the intracranial vessels. The pulse sequence used typically consists of a gradient recalled echo (GRE) sequence, which is acquired with either a two-dimensional (2D) or 3D technique. 2D TOF images are obtained as contiguous or slightly overlapped sections, whereas 3D TOF images are derived from one or more overlapping 3D volumes. Of these two methods, 3D TOF is more commonly used than 2D TOF because of its superior spatial resolution.²⁴

In PC MRA, the vascular contrast is obtained by applying a bipolar phase-encoding gradient and a velocity-encoding (VEnc) factor.^25,²⁶ Phase shifts in moving spins or flowing blood are obtained when gradients with opposing polarities are applied twice during a single RF excitation. This method results in nulling of signal from stationary tissue while exploiting signal from flowing blood. Visualization of blood flow depends on the velocity of the flow. In specific encoding orientations, middle-gray usually represents no flow, whereas progressively increasing shades of white or black indicate directionally encoded higher velocities up to the predefined VEnc factor. Blood that flows at velocities higher than the VEnc factor will demonstrate aliasing, a situation in which signal intensity “wraps around” and will therefore be opposite that expected. In other words, if flow in a certain direction matches the VEnc factor, it will be encoded as white, but any greater velocity in that direction will actually appear as black. Because arterial flow is more rapid than venous flow, the typical arterial VEnc factor intracranially measures greater than 60 to 80 cm/sec, whereas slow flow within veins and venous sinuses is best imaged with a VEnc factor of 20 cm/sec or less.²⁶ Images generated from this technique provide not only direction of flow but also magnitude of flow.

CE MRA uses a high-resolution T1-weighted technique in which the vessels are accentuated by a bolus of gadolinium chelate. The 2% to 3% concentration of gadolinium contrast agent in blood (by volume) causes a marked T1-shortening effect in vessels with a resultant increase in vessel signal. CE MRA provides anatomic information related to lumen diameter and the concentration of contrast material in vessels rather than physiologic information reflecting the flow rate, as calculated with TOF and PC techniques. Advantages of CE MRA over non-CE techniques include higher signal-to-noise ratios, decreased susceptibility to artifacts caused by pulsatility and flow, shorter image acquisition times, and temporal resolution that illustrates patterns of flow over time. Older CE MRA paradigms (e.g., elliptic-centric techniques) had collected full RF spectral data during the bolus, so only about three full-volume MRA repetitions could be acquired per minute. However, with that technique, the intended arterial-phase imaging was frequently obscured by venous enhancement. The need for improved temporal resolution led to the development of time-resolved CE MRA, such as TRICKS (time-resolved imaging of contrast kinetics) and TREAT (time-resolved echo-shared angiography technique).²⁴ These sequences improve temporal resolution by repetitively acquiring the center of the RF spectrum along with varying peripheral portions during each acquisition; data missing from unsampled regions are interpolated by using data obtained at other points in time. This development enables more rapid 3D acquisitions before, during, and after transit of a contrast bolus, thereby allowing identification of the arterial, capillary, and venous phases.^27,²⁸

Clinical Uses and Applications

The angiographic modality with unparalleled spatial and temporal resolution is DSA. However, manipulation of catheters in the aorta or in the carotid or vertebral arteries, even by an experienced operator, is associated with a 1.3% risk for a cerebral event.²⁹ Therefore, in clinical practice most institutions still use noninvasive techniques such as computed tomographic angiography (CTA) or MRA for preliminary evaluation of the vascular tree in the clinical setting of infarction, aneurysm, AVM, dural arteriovenous fistula (AVF), and veno-occlusive disease.

Stroke

A cerebrovascular accident, or stroke, is major cause of morbidity and mortality. It is the third most common cause of death in the United States, with approximately 795,000 cases occurring annually. The majority of these cases are caused by atherosclerotic disease, which results in stenosis and progressive occlusion of an arterial vessel. This can lead to ischemia and potential infarction of the brain tissue involved. MRA has been used to study the major arterial intracranial vessels to evaluate the effect of atherosclerotic disease. Specifically, 3D TOF MRA has been used as a screening method in stroke patients because it is minimally invasive and provides a reasonable assessment of the degree of vessel occlusion. Studies have shown 100% detection of vessel occlusion on 3D TOF MRA when correlated with DSA; however, grading of stenosis is less accurate (61%) when using MRA.³⁰ Additionally, visualization of vessels that demonstrate slow or turbulent flow can be improved with the use of contrast material injected intravenously. For example, CE MRA has been shown to offer better visualization of the internal cerebral and middle cerebral arteries that exhibit artifactual narrowing on 3D TOF MRA because of slow or turbulent flow.³¹ Although such limitations exist, TOF MRA remains an important sequence for evaluation of the intracranial arterial circulation in patients with acute neurological symptoms.

Intracranial Aneurysms

Detection of intracranial aneurysm can be done with noninvasive techniques such as MRA and CTA. 3D TOF MRA has been shown to have an overall sensitivity of 87% and specificity of 95% for the detection of intracranial aneurysms (Fig. 19-9). The sensitivity for detection of aneurysms larger than 3 mm is greater than that for aneurysms 3 mm or smaller (94% versus 38%) on older 1.5T scanners.³² At a magnetic field strength of 3T, aneurysms as small as 1 mm can be detected.³³ Both TOF MRA and CE MRA are optimized to show vessel lumens; apical thrombus may remain undetected unless MRA is correlated with conventional sequences that are more sensitive for nonflowing blood.