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1 Both authors: Department of Radiology, Duke University Medical Center, Box 3808, Durham, NC 27710.
Received July 27, 1999;
accepted after revision December 15, 1999.
Address correspondence to J. R. Petrella
Introduction
 Top Introduction TechniquesClinical ApplicationsConclusionReferences |
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To obtain hemodynamic parameters from serial tissue tracer concentration measurements, a general model of the method by which that tracer passes through or distributes in the target organ is required [4]. Such a model must be based on an understanding of the manner in which the tracer is infused—that is, bolus injection versus constant infusion, and on assumptions about the pharmacokinetic properties of the tracer in the organ of interest. These assumptions include diffusibility from the intravascular to extravascular space, volume of distribution, and equilibrium half-life of the tracer.
Exogenous Tracer Methods
Exogenous tracer methods in MR perfusion imaging use a model that assumes
the tracer is restricted to the intravascular compartment and does not diffuse
into the extracellular space. Imaging can be performed either dynamically
(rapid imaging over time during a bolus injection) or in the steady state
(imaging after a constant infusion has reached an equilibrium concentration in
the blood).
Dynamic imaging.—Dynamic imaging takes advantage of transient changes in the local magnetic field of the surrounding tissue induced by a bolus of paramagnetic tracer passing through the organ capillary network (Fig. 1A,1B,1C,1D,1E). These changes in the local magnetic field can be measured as signal changes on MR imaging. Ultrafast imaging techniques, such as echoplanar and spiral MR imaging [5, 6], enable the accurate measurement of rapidly varying signal changes that are due to the first pass of the bolus with adequate temporal resolution (<2 sec for coverage of the entire brain). Signal-time course data can then be converted to relative tracer tissue concentration-time course data [3]. Tracer concentration-time curves can then be analyzed to determine various tissue hemodynamic parameters, such as tissue blood volume, blood flow, transit time, and bolus arrival time (Fig. 2). In this article, the terms "cerebral blood volume," "cerebral blood flow," and "mean transit time" are defined as follows: Cerebral blood volume refers to the volume of blood in a given region of brain tissue, commonly measured in milliliters per 100 g of brain tissue. Cerebral blood flow refers to the volume of blood per unit time passing through a given region of brain tissue, commonly measured in milliliter per minute per 100 g of brain tissue. Mean transit time refers to the average time it takes blood to pass through a given region of brain tissue, commonly measured in seconds. Bolus arrival time refers to the time it takes for an IV-injected bolus of contrast material to arrive at a given region of the brain, also commonly measured in seconds.
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These parameters are dependent on the specific features of the bolus injection, including the amount of contrast material injected, the injection rate, and the paramagnetic properties of the contrast agent. In addition, hemodynamic parameters depend on variables within the subject being imaged, such as total-body vascular volume and cardiac output. As a result, hemodynamic parameters cannot be directly compared between different subjects and may even differ between examinations on the same subject at different times. Nonetheless, semiquantitative or relative values can be obtained using an internal standard of reference such as normal-appearing gray or white matter. This allows intra- and intersubject comparisons. Semiquantitation is valuable for subcategorizing disease processes, following disease course, and monitoring the effects of therapeutic interventions, provided such alterations are local and do not change the hemodynamic parameters of the internal reference. For example, in the case of stroke, relative hemodynamic measurements can be useful in monitoring the periinfarct ischemic tissue zone at risk for infarction as well as in monitoring the effects of thrombolytic therapy [7, 8]. For diffuse disease processes, in which the internal reference may also be affected, absolute quantitation is necessary. Although absolute quantitation of cerebral blood volume and cerebral blood flow has been attempted using dynamic MR methods that measure arterial input to the brain [9,10,11,12,13], the accuracy of these methods remains unproven [14]. With further improvements in signal-to-noise capability and improved techniques to determine true arterial input, quantitation of absolute cerebral blood volume and cerebral blood flow with dynamic MR methods may well be feasible.
Determination of relative cerebral blood volume from tracer concentration-time data is straightforward and robust, accomplished by integrating the area under the tracer concentration-time curve (Fig. 2). This integration may be performed on the curve data points themselves or on an analytic fit of the data points [15]. The latter approach has the advantage of eliminating overestimation from the effects of tracer recirculation, but this approach has the disadvantage of requiring high signal stability and faster imaging over time [16]. Determination of relative cerebral blood flow requires more extensive processing of the imaging data and is more adversely influenced by poor image quality and instability in the MR signal over time. The processing techniques require deconvolution of an arterial input function from tissue concentration-time data to find the true brain clearance, or mean transit time through the cerebral capillary bed (mean transit time). Cerebral blood volume, calculated by integrating the area under the deconvolved tissue concentration-time curve, is then divided by mean transit time to obtain cerebral blood flow [17]. Alternatively, the initial height of the deconvolved tissue concentration-time curve may be taken as the cerebral blood flow, and the mean transit time may then be calculated as the ratio of cerebral blood volume to cerebral blood flow [18] (Fig. 2). Again, image quality and signal stability over time are important requirements for reliably calculating relative cerebral blood flow because the deconvolution technique mentioned previously can amplify noise and artifactually introduce bias [18].
Determination of an accurate arterial input function is also an important requirement for calculating relative cerebral blood flow. An arterial input function may be obtained directly from the imaging data by manually selecting the voxels from which the arterial input function will be obtained [19]. This may be aided by narrowing the selection to a small population of voxels chosen using an automated algorithm that searches the entire imaging volume for voxels with time-concentration curves that satisfy criteria characteristic of arteries, such as a large peak, early arrival time, and a short mean transit time [9, 11]. The use of such an algorithm increases reproducibility because it requires less user interaction. It should be noted, however, that the exclusive reliance on an automated approach may lead to erroneous selection of an arterial input function. For example, in the case in which a cerebral hemisphere is being fed by a diseased middle cerebral artery, deriving the arterial input function from voxels in the diseased hemisphere would lead to a more accurate result than deriving the arterial input function from the voxels in the contralateral hemisphere, even though the latter may better satisfy criteria for a "normal" artery. Voxel-by-voxel determination of cerebral blood flow, in theory, requires determination of the unique arterial input to each voxel. Because this is not possible, most methods assume the arterial input is uniform across the brain and apply a single arterial input function to the entire brain. In the case of a unilaterally diseased middle cerebral artery, this assumption is violated. Applying an arterial input function chosen from voxels in the normal contralateral hemisphere may lead to underestimation of cerebral blood flow and false-positive identification of an ischemic zone.
Dynamic sequences must be ultrafast to monitor the rapid first-pass transit of a bolus of contrast agent through the brain, which is on the order of 18 sec [20]. Either T1- or T2-weighted techniques can be used. The T2-weighted sequences are more commonly used in clinical practice. Using these sequences, injection of a paramagnetic contrast agent causes a transient drop in signal intensity that is due to the susceptibility effects of the paramagnetic contrast agent. A single- or dual-slice dynamic study can be performed on a conventional MR scanner without specialized gradient hardware [9]. Multislice techniques (up to 30 slices per second) are available on systems with specialized gradient hardware for echoplanar imaging or spiral imaging [21]. These techniques can be either T2-weighted (spin echo) or T2*-weighted (gradient echo). The spin-echo technique has the advantage of minimizing artifact at brain-bone and brain-air interfaces and is more sensitive to signal changes from paramagnetic contrast material passing through small vessels, such as capillaries, rather than through large vessels, such as cortical veins [22]. The spin-echo technique has the disadvantage of requiring a larger dose of contrast material, often 1.5-2.0 times that of a standard dose, to produce signal changes equivalent to those of the gradient-echo technique (Aronen H et al., presented at the Society of Magnetic Resonance in Medicine meeting, August 1992). Furthermore, the spin-echo technique may create bias on serial studies, leading to artificially elevated cerebral blood volume measurements, if repeated within 2 hr of the initial study. This bias is caused by a residual contrast material effect that alters the magnitude of signal change from baseline [23]. These effects have been shown to not be significant using a gradient-echo technique [11].
A T1-weighted dynamic technique is another method by which to measure cerebral hemodynamics [24] and has the advantage of requiring a smaller contrast material dose and providing better temporal resolution than the T2- or T2*-weighted sequences. The T1-weighted technique measures the relaxivity effects, rather than the susceptibility effects, of an IV-injected dose of paramagnetic contrast material. The relaxivity effect of paramagnetic contrast material refers to the shortening of T1 relaxation time, leading to higher signal on T1-weighted images, whereas the susceptibility effect refers to the shortening of T2 and T2* relaxation times, leading to lower signal on T2- or T2*-weighted images. Because the relaxivity effects of gadopentetate dimeglumine are much stronger than the susceptibility effects, the T1-weighted pulse sequences require a smaller amount of contrast material (approximately 10%) than the T2- or T2*-weighted techniques [25], allowing multiple repeated studies. Moreover, the short injection time allowed by a smaller bolus may result in better quantitation of cerebral blood volume and cerebral blood flow provided that the temporal resolution of the pulse sequence allows tracking the bolus over a sufficient number of time points to extract the corresponding parameters [24]. Subsecond imaging times (300-900 msec) over an anatomic range of one to two slices are currently possible with this technique using fast T1-weighted gradient-echo imaging [25]. The T2- or T2*-weighted technique requires imaging times on the order of 1.5-2 sec, although the anatomic coverage is greater with echoplanar imaging (8-11 slices) or spiral imaging (18-20 slices) [21, 26]. The disadvantage of the T1-weighted technique is that leakage through the blood-brain barrier may lead to errors in measurements of hemodynamic parameters. Although this may be corrected for in the calculations, the effects of blood-brain barrier breakdown are greater with the T1-weighted technique than with the T2- or T2*-weighted technique. Quantitative assessment of permeability through the blood-brain barrier has been examined using both T1- and T2-weighted techniques. Further discussion of such techniques can be found in the literature [27, 28].
Steady-state imaging.—In addition to the more commonly used T1- and T2- or T2*-weighted dynamic perfusion imaging techniques, a T1-weighted steady-state technique may be used to estimate absolute cerebral blood volume with high spatial resolution across the entire brain. This method assumes the tracer is nondiffusible from the intravascular to extravascular space. With this technique, a baseline image is obtained before the injection of the paramagnetic agent, followed by a postinfusion image that is acquired during the "steady state,"—that is, up to 30 min after the contrast material has circulated through the body and reached a point of relative concentration equilibrium. By subtracting the baseline image from the postcontrast steady-state image and normalizing the pixel values to those of a pixel containing only blood, such as in the sagittal sinus, one can obtain a map of absolute cerebral blood volume in units of volume percent [29, 30]. This can be converted to the more conventional units of milliliters per 100 g of brain by normalizing to the density of brain tissue (1.04 g/ml) and multiplying by 100 [9]. Unfortunately, this approach has a number of disadvantages. First, because image subtraction is performed, the resulting images have a low signal-to-noise ratio. Second, patient movement between the pre- and postcontrast scans may affect the accuracy of these measurements. Third, spurious results can be obtained in areas where the blood-brain barrier has been disrupted and the assumption of tracer nondiffusibility has been violated; therefore, the technique is not useful in many cases of infarction and tumor.
Endogenous Tracer Methods
Endogenous tracer methods in MR perfusion imaging use a model that assumes
the tracer freely diffuses from the intravascular compartment into the tissue
compartment. This model is similar to that used in positron emission
tomography and single-photon emission computed tomography (SPECT) in which a
tracer is administered and the regional accumulation, influenced by regional
blood flow and tracer half-life, is measured
[31]. Endogenous tracer MR
perfusion methods take advantage of signal loss resulting from magnetically
labeled water protons (spins) flowing into the imaging plane and exchanging
with tissue protons. Water protons within inflowing arterial blood are
magnetically labeled (or "tagged") by the application of a special
radiofrequency pulse designed to invert spins in a thick slab proximal to the
slice of interest (Fig. 3). By
measuring signal changes between tagged images and baseline untagged images,
qualitative or quantitative images of cerebral blood flow can be obtained
(Fig. 4). Inflowing blood may
be tagged continuously or intermittently
[31,
32]. Although
continuous-labeling techniques afford twice as much signal contrast compared
with pulsed techniques, they produce substantially more radiofrequency
pulse-induced power deposition to the subject. This safety consideration can
ultimately limit slice coverage and acquisition time.
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Spin-tagging techniques suffer inaccuracies that are due to two major effects. The first is magnetization transfer effects in the imaging plane as spins are labeled below the imaging plane with a radiofrequency pulse. During this time, the spins in the plane of interest experience an off-resonance radiofrequency pulse that selectively saturates the broad resonance peak of macromolecular-bound protons. This saturation is then transferred to the free-proton pool, resulting in as much as a 60% loss in observed brain signal [33]. To compensate for this effect, one strategy involves applying a radiofrequency pulse during the baseline state at an equal distance above the plane of interest [34, 35].
The second inaccuracy is from the loss of spin labeling during the arterial transit period due to T1 relaxation as blood moves from the tagging plane to the imaging plane [31]. These arterial transit effects may be markedly reduced by introducing a delay after continuous labeling [36] or by tagging spins immediately below the imaging plane using an intermittent pulse to reduce overall transit time [32]. The latter technique, called "EPISTAR" (echoplanar imaging with signal targeting and alternating radiofrequency), provides only a qualitative map of cerebral blood flow because the relationship between cerebral blood flow and EPISTAR signal is complex and depends on differential arterial transit times, the angle of feeding arteries with the imaging plane, and inflow effects [37]. This technique also suffers from low sensitivity and therefore low flow rates (10-25 ml · 100 g-1 · min-1) may be difficult to detect.
An alternative to labeling spins proximal to the imaging plane is to directly label spins in the imaging plane itself using a slice-selective inversion-recovery technique and thereafter measure signal increases from unlabeled inflowing spins. In this case, the unlabeled spins have complete longitudinal magnetization and the T1-relaxation effects that are due to arterial transit are eliminated. Thus, signal changes are indirectly related to cerebral blood flow. By applying an alternating global inversion-recovery pulse along with the slice-selective pulse and comparing the two conditions, one can measure signal changes that are caused solely by inflowing blood. These signal changes are more directly related to absolute cerebral blood flow. The FAIR (flow-sensitive alternating inversion-recovery) technique [38] is one example of this method. A multislice version of FAIR has recently been developed using imaging [39].
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MR perfusion imaging is useful not only in the assessment of stroke, but also in the assessment of stroke risk. Under normal circumstances, the brain has an autoregulatory mechanism for maintaining adequate cerebral oxygenation in the face of decreasing cerebral perfusion pressure, which allows normal blood flow despite fluctuations in systemic pressure. This mechanism may be impeded in patients with hemodynamically significant carotid artery stenosis who are at high risk for stroke. The ability to maintain an autoregulatory response to hemodynamic stress has been termed "cerebrovascular reserve capacity." Areas of the brain supplied by a markedly stenotic or occluded artery, in which vasodilatation has already occurred to maintain adequate flow, lack cerebrovascular reserve capacity. As a result, when a pharmacologic vasodilatory challenge is administered, minimal vasodilatory response occurs. Assessment of response to vasodilatory challenge has therefore also been used as an indirect means of measuring cerebrovascular reserve capacity. Perfusion MR imaging may be used to assess cerebral blood volume or cerebral blood flow before and after a vasodilatory challenge using agents such as carbon dioxide or the carbonic anhydrase inhibitor, acetazolamide [10,11,12]. In addition to a poor response to a vasodilatory challenge, perfusion imaging may show other abnormalities in the cerebral hemisphere ipsilateral to a severe carotid stenosis or occlusion, such as delayed bolus arrival time and prolonged mean transit time [42, 43] (Fig. 6A,6B,6C,6D).
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Brain tumors.—Another area in which MR perfusion imaging may be useful is in the evaluation of brain tumors. Dynamic imaging is performed using either T1-weighted or T2- or T2*-weighted technique [44]. Cerebral blood volume maps can be used to assess neovascularity in tumors, which is thought to correlate with tumor grade and malignant histology. Because of selective sensitivity to small vessels, the T2-weighted technique may be preferred over the T2*-weighted technique [22]. Cerebral blood volume maps may aid in early evaluation of therapeutic agents, especially of a new class of drugs aimed at suppressing growth of tumor blood vessels. These maps can also potentially be used to localize areas of tumor more likely to yield positive results on stereotactic biopsy and to noninvasively differentiate radiation necrosis from recurrent tumor in circumstances in which conventional MR findings are equivocal [45, 46] (Figs. 7A,7B,7C and 8A,8B,8C). Similar techniques have evolved for differentiating radiation necrosis from recurrent tumor on the basis of differences in blood-brain barrier permeability. For detailed discussions of these methods, the reader is referred to the literature [28, 47].
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Other disorders.—In addition to evaluation of ischemia and tumors, MR perfusion imaging has been applied to the study of various other neurologic and psychiatric disorders, such as dementia and migraine headaches [48]. The effects of psychoactive drugs, such as cocaine, have been studied as well (Kaufman MJ et al., presented at the International Society of Magnetic Resonance in Medicine, April 1996). In the case of migraine headaches, decreases in cerebral blood volume and cerebral blood flow have been seen during the auras compared with the post-aura state (Sorensen AG et al., presented at the International Society of Magnetic Resonance in Medicine, April 1996). In the case of dementia, decreases in cerebral blood volume in the temporal and parietal lobes of patients with Alzheimer's disease have correlated well with the results of SPECT studies on the same subjects [49, 50].
Endogenous Tracer Methods
Clinical applications of the endogenous tracer methods have been limited
compared with those of exogenous methods because of longer acquisition times
and sensitivity to patient motion. Furthermore, although FAIR and other
spin-labeling techniques are capable of calculating absolute cerebral blood
flow in theory, in practice these techniques have not yet shown sufficient
signal-to-noise ratio to validate these measurements, especially in low-flow
states and in white matter. In such situations, even the theoretical
assumptions required to make absolute flow estimates may break down because
of, for example, the loss of spin labeling from prolonged transit times. Once
these limitations are overcome, however, many clinical applications may be
possible.
Stroke.—Quantitative MR cerebral blood flow measurements could potentially be obtained as part of a complete MR evaluation of stroke in the assessment of tissue viability and stroke etiology, for example. Previous work has suggested that tissue viability, in the setting of acute stroke, is related to the degree and duration of ischemia [51, 52]. The degree of ischemia has been assessed in the past through quantitative regional cerebral blood flow measurements using positron emission tomography and 133Xe SPECT, and thresholds of cerebral blood flow measurements have been established below which tissue viability is unlikely. The role of hypoperfusion as the primary cause of stroke can also be examined using quantitative MR measurements of cerebral blood flow in patients with potential border-zone or watershed infarctions [53].
Clinical research.—Another major benefit of absolute quantitation of cerebral blood flow is in clinical research, which depends on accurate inter- and intrasubject comparisons. Intrasubject comparisons are useful in following the natural history of a disease process or in assessing the effect of therapeutic interventions in conditions such as stroke, neoplasms, and neurocognitive disorders such as dementia, especially in cases when no normal areas of internal reference are available. Intersubject comparisons of cerebral blood flow between different patient populations may allow assessment of drug efficacy or may be a useful tool in investigating disease mechanisms.
Functional brain mapping.—A third major benefit of absolute quantification of cerebral blood flow is in the area of functional MR brain mapping. Cerebral activation is usually determined through qualitative assessment of local changes in deoxyhemoglobin concentration. This phenomenon, in which local differences in cerebral blood oxygenation levels are related to the degree of neuronal activity, is known as the blood oxygenation level-dependent (BOLD) effect. The BOLD effect is based on close coupling of neuronal activity to hemodynamic response in the brain [54]. Although BOLD changes are more sensitive to cerebral activation than absolute cerebral blood flow changes, BOLD changes are dependent on a number of other physiologic variables and MR parameters [38, 55]. Direct measurement of cerebral blood flow changes during cerebral activation may enable better localization of neuronal activity than measurement of BOLD changes and may allow more physiologically meaningful comparison of activation between different brain regions [38] (Fig. 9A,9B). Quantitative cerebral blood flow changes during motor and working memory tasks have been measured with the arterial spintagging technique and yield results similar to those obtained with other techniques such as 133Xe SPECT and positron emission tomography [56, 57].
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Acknowledgments
We thank Joseph Frank and Alan McLaughlin at the National Institutes of
Health for their helpful comments and Jimmie Wong and Luiz Celso H. Cruz for
assistance in preparing the figures.
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