To quantitate cerebral blood flow (CBF) in the entire brain using the 3D echo planar imaging (EPI) PULSAR (pulsed star labeling) technique.
To quantitate cerebral blood flow (CBF) in the entire brain using the 3D echo planar imaging (EPI) PULSAR (pulsed star labeling) technique.
The PULSAR technique was modified to 1) incorporate a nonselective inversion pulse to suppress background signal; 2) to use 3D EPI acquisition; and 3) to modulate flip angle in such a manner as to minimize the blurring resulting from T1 modulation along the slice encoding direction. Computation of CBF was performed using the general kinetic model (GKM). In a series of healthy volunteers (n = 12), we first investigated the effects of introducing an inversion pulse on the measured value of CBF and on the temporal stability of the perfusion signal. Next we investigated the effect of flip angle modulation on the spatial blurring of the perfusion signal. Finally, we evaluated the repeatability of the CBF measurements, including the influence of the measurement of arterial blood magnetization (a calibration factor for the GKM).
The sequence provides sufficient perfusion signal to achieve whole brain coverage in ≈5 minutes. Introduction of the inversion pulse for background suppression did not significantly affect computed CBF values, but did reduce the fluctuation in the perfusion signal. Flip angle modulation reduced blurring, resulting in higher estimates of gray matter (GM) CBF and lower estimates of white matter (WM) CBF. The repeatability study showed that measurement of arterial blood signal did not result in significantly higher error in the perfusion measurement.
Improvements in acquisition and sequence preparation presented here allow for better quantification and localization of perfusion signal, allowing for accurate whole-brain CBF measurements in 5 minutes. J. Magn. Reson. Imaging 2011;33:287–295. © 2011 Wiley-Liss, Inc.
ARTERIAL SPIN LABELING (ASL) (1) provides noninvasive measurements of cerebral blood flow (CBF) by using labeled blood as an endogenous contrast agent. Measurement of cerebral perfusion can provide unique information not achievable with other magnetic resonance imaging (MRI) contrasts in a number of clinical situations, eg, differentiation of tumor recurrence from radiation necrosis. Clinically, bolus tracking of an exogenous gadolinium-based contrast agent (GBCA) is typically performed using dynamic susceptibility imaging. Use of tagged blood as contrast has several advantages as compared with techniques using exogenous dynamic contrast agents. Foremost of these is the elimination of risks associated with the administration of GBCA including anaphylaxis and nephrogenic systemic fibrosis. Expense and dose limitations also preclude multiple repetitions of the GBCA perfusion experiment.
Commercial implementations of ASL perfusion MRI are currently based on 2D multislice acquisitions. These sequences provide limited spatial coverage of the brain. Clinically there is a need for methods that provide whole-brain coverage in a reasonable time, motivating the use of 3D acquisition.
Three-dimensional acquisition has a number of potential benefits over 2D acquisition. In addition to permitting whole-brain coverage and possibly higher signal-to-noise ratio (SNR), 3D acquisition eliminates slice-dependent variation of the perfusion signal resulting from slice-dependent acquisition delay times in 2D acquisition. This is due to the fact that the SNR of 3D images is determined by the signal energy at the (kx, ky, kz) = 0 point. The acquisition delay time can then be considered to be from the application of tagging or control pulses to the echo time of the acquisition through the center of k-space. Three-dimensional acquisition of the imaging volume without gaps can also provide better reproducibility in longitudinal ASL-based functional (f)MRI studies. For these reasons, a number of investigators have made use of 3D acquisition, either as segmented or single-shot acquisitions. Examples of segmented 3D ASL techniques include the use of multishot spirals with pulsed arterial spin labeling (2) and with continuous arterial spin labeling (3). Single-shot techniques reduce variation in signal arising from phase inconsistencies in multishot techniques. 3D gradient- and spin-echo (GRASE) acquisition was implemented with flow-sensitive inversion recovery (FAIR) pulsed arterial spin labeling (4) and with continuous ASL (CASL) (5). The use of single-shot 3D fast spin-echo spiral acquisitions has also been demonstrated (6). While fast spin-echo and GRASE acquisitions have the benefit of reducing T2* related artifacts that occur with gradient-echo echo planar imaging (EPI) acquisitions, this comes at the expense of higher radiofrequency (RF) power deposition and prolonged acquisition times relative to pure EPI readouts. Talagala et al (7) demonstrated the use of nonsegmented 3D EPI with CASL in which a complete 3D dataset was acquired using a train of z-phase encoded, low flip angle 3D EPI readouts following each CASL preparation interval. This approach, referred to here as 3D turbo field echo planar imaging (TFEPI) (turbo refers to the acquisition under nonsteady-state conditions), provides high k-space sampling efficiency. With centric encoded 3D TFEPI, true whole-brain coverage using up to 35 slices with 4-mm slice thickness can be obtained. In addition, for high field imaging, SAR can be reduced by reducing the excitation angle used with 3D TFEPI while still maintaining sufficient SNR. Therefore, in this work we further develop the use of 3D TEFPI for ASL perfusion imaging by combining it with the pulsed star labeling (PULSAR) technique (based on the EPISTAR technique (8)) described by Golay et al (9, 10). Here we investigate modifications to the basic sequence to improve the temporal stability of the perfusion images, reduce blurring in 3D TFEPI images, and improve quantitation and reproducibility of perfusion data.
Previously, it has been shown with 3D segmented acquisition FAIR (2) that use of inversion pulses results in superior temporal stability of perfusion images by reducing the background signal. Background suppression pulses have also been used in single-shot 3D ASL studies (4–6). With 3D TFEPI, use of a centric-ordered acquisition can increase the effectiveness of background suppression pulses. However, use of multiple background suppression pulses can cause significant attenuation of the ASL signal (2, 11), which can offset the advantages of background suppression. Accordingly, we employ a single nonselective inversion pulse (hyperbolic-secant pulse) to reduce temporal variability of the perfusion signal. Quantitative CBF values are compared with and without background suppression in gray matter and globally.
One potential complication with 3D TFEPI readout arises from the application of constant flip angles to acquire z-phase encodes under nonsteady-state conditions. This results in blurring along z-axis due to modulation of kz space, and hence, in errors in CBF values of gray matter (GM) and white matter (WM). We analyze the effect of nonsteady-state constant flip angle imaging and correct for blurring by modulating the flip angle train such that magnetization in GM remains almost constant across all 3D-TFEPI shots.
To derive quantitative values, a general kinetic model (GKM) (12) is favored for pulsed arterial techniques due to its simplicity and ease of application. Application of the model requires 1) a well-defined bolus of tagged blood, and 2) complete delivery of the entire bolus to the voxel imaged at the time of measurement. In this work a well-defined bolus was obtained by incorporating a QUIPSS II saturation pulse (13) into the PULSAR sequence.
The most important measurement factor affecting CBF quantification with the GKM model is the equilibrium magnetization of the arterial blood signal. We therefore performed a repeatability study in order to quantify the error resulting from measurement of this factor.
The original PULSAR sequence (9) was modified to include a QUIPSS II saturation pulse (13) and an inversion recovery (IR) pulse as shown schematically in Fig. 1. The preparation scheme consists of a four pulse presaturation sequence to saturate magnetization in the imaging slab. This is followed by the labeling pulse, which consists of one 180° adiabatic hyperbolic secant inversion pulse. For the case of the control sequence, two adiabatic inversion pulses with the same time duration and bandwidth as the labeling pulse are used. The amplitude of both inversion pulses for the control sequence is lower than the labeling pulse. The labeling pulse inverts spins in the selected labeling slab (below the imaging slab), while the control pulse does not result in any net magnetization. However, the control pulse compensates for the magnetization transfer effect. Introduction of a QUIPSS II saturation pulse helps define bolus length and the nonselective IR pulse helps suppress the signal in GM, WM, and cerebrospinal fluid (CSF). This is followed by the imaging slab acquisition, which consists of a train of low flip angle gradient echo planar imaging sequences with phase encoding along the z (slice) direction (7, 14). We refer to the PULSAR sequence with 3D TFEPI acquisition (but no inversion pulse) as 3D PULSAR and that with the adiabatic inversion pulse is referred to as IR-3D-PULSAR.
Quantification of CBF was done using f(TD) = ΔM/[2ηM0Aτ exp(–TD/T1A)] (eq.  in (12)), where ΔM is the perfusion signal, η is the inversion efficiency, M0A is the equilibrium magnetization of arterial blood, τ is the duration of the bolus, TD is the delay between tagging and acquisition, and T1A is the T1 of arterial blood. T1A is assumed to be 1650 msec at 3T (15). The inversion efficiency, η, is assumed to be 0.91 (derived on a Philips 3T scanner, Best, Netherlands (10)). For centric-ordered 3D acquisition, TD is defined by the time between the tagging inversion pulse and the kz = 0 slice encoding, which for our centric-ordered case corresponds to approximately the beginning of data acquisition. M0A is difficult to measure directly, as intracranial arteries are small and subject to partial volume errors. Therefore, we chose to measure M0V directly from a region of interest (ROI) in the sagittal sinus and convert this to M0A by correcting for the different transverse relaxation rates of arterial (R2A*) and venous (R2V*) blood by the equation M0A = M0V · exp [TE · (R2V* – R2A*)]. The relaxation rate of arterial blood, R2V*, was assumed to be 46.2 s−1 and R2A* = 18.8 s−1 at 3T (10).
Measurement of blood signal in the sagittal sinus is done on the first phase of the 3D-PULSAR image set to ensure that the signal in the sagittal sinus is free of any prepared (inversion or saturation) magnetization, as the mean transit time through the vascular bed is ≈4 seconds (12) (much greater than the TD time of 1.8 second used here). A separate sequence, which in turn entails extra processing due to differences in image scaling, was therefore not required.
For the case when an IR pulse is used (IR-3D-PULSAR), an extra inversion efficiency factor is applied to M0A (=M0A · η) derived from the first phase of the 3D-PULSAR images in addition to image scaling factors. For the IR-3D-PULSAR acquisition, ΔM was calculated as (Label – Control) instead of (Control – Label) images to account for the inversion pulse (2).
A single inversion pulse was introduced with inversion time around 900 msec. In combination with the saturation of the imaging slab, this results in GM, WM, and CSF longitudinal magnetization values of 0.25, 0.43, and 0.03 (based on published values for T1/T2) for centric encoded data acquisition at 1800 msec after labeling and for the repetition time employed during scanning. This is in contrast to corresponding values of 0.75, 0.88, and 0.33 when no inversion pulse is used. Consequently, background signal is suppressed in all three tissue types.
The acquired ΔM(t) images were analyzed for temporal stability of the perfusion signal by calculating 〈σ(ΔM(t))〉/〈ΔM(t)〉 globally for the whole brain (Glo) and for ROIs in GM, WM, and CSF. All ROIs and slices were matched for both sequences.
In addition, CBF estimates in GM were compared for the two sequences. Erosion and dilation (to remove skull) followed by automated segmentation based on Otsu's algorithm (16) available in MatLab (MathWorks, Natick, MA) was applied to all CBF maps to separate regions of high perfusion (approximating GM) from lower perfusion (approximating WM). Mean GM CBF value was determined by averaging values over all slices.
Typically, 3D acquisition under nonsteady-state conditions results in blurring. For ASL techniques, this reduces GM CBF and increases WM CBF values due to partial volume effects. In 3D-TFEPI, the longitudinal magnetization after each RF excitation pulse can be described by (17):
and the transverse magnetization after each RF pulse is:
where E1 = exp(–TR/T1), E2* = exp(–TE/T2*), α is the excitation angle, and L is the total number of slice encodings. Use of a constant flip angle results in a decay of the transverse magnetization towards its final steady-state value. This signal modulation (which appears along slice phase encoding in 3D TFEPI acquisition) is the cause of the blurring. To correct for this modulation, the flip angles can be adjusted such that the magnetization stays approximately constant across all slice encodings. This is achieved if the flip angle is calculated as (17, 18):
A particular final flip angle can be targeted and the flip angle schedule calculated within a few iterations (19). Figure 2a shows one such flip angle train calculated using GM T1/T2 values (1300 msec and 60 msec, respectively), TR/TE = 20/9 msec (corresponds approximately to TR/TE times used on the scanner), a final flip angle of 30°, and with 31 slice encodings. (This corresponds to 24 slice acquisitions with oversampling typically used with 3D acquisition.) Figure 2b shows the transverse magnetization for the constant flip angle (30°) case and for the modulated flip train. Magnetization remains mostly constant for the modulated flip angle train. The point spread function (PSF) of Mxy(n) (Fig. 2c) clearly shows that increased blurring results from using the constant flip angle train. Figure 2D shows the effect of the PSF on a slice of thickness 4 mm. Constant flip angle scheme results in increased smearing of the signal along the slice direction. The peak signal (normalized to the ideal signal) is just 0.52, while it is 0.98 with the variable flip angle scheme. Simulations show that transverse magnetization (and hence SNR) increases with increasing final flip angle. However, increasing the final flip angle from 30° to 90° results in a mere 3% increase in magnetization at the center of k-space. In addition, on our scanner, increasing the final flip angle results in increasing values for TE/TR of the single-shot EPI acquisition due to RF pulse stretching. As a result, a final flip angle of 30° was used with variable flip angle excitation and for 31 slice encodings.
The derived CBF values follow an inverse relationship with arterial blood signal (M0A). As described earlier, measurement for M0A is typically done in a voxel in the sagittal sinus that is completely filled with blood and free of partial volume effects. Measurement of M0A is considered to be an important source of error in CBF determination using GKM with a bolus cutoff pulse. We investigated the repeatability of CBF values across volunteers with and without this potential source of error (20).
For this purpose, four 3D PULSAR and four IR-3D-PULSAR scans were performed on each volunteer in the same session. M0A was derived from signal measured in the sagittal sinus for each 3D PULSAR scan. GM CBF values were calculated for each case using the corresponding measured M0A value. CBF values were also calculated for the four scans using an average value for M0A to reflect other (pulse imperfections, differences in transit delay etc) errors. Since introduction of an IR pulse alters signal in the sagittal sinus, the average value of M0A (denoted 〈M0A〉) from the 3D PULSAR scan was used for CBF calculation for IR-3D-PULSAR. While CBF values obtained from 3D-PULSAR scans reflect variation mainly due to M0A, values obtained using 〈M0A〉 reflect other errors. The mean and standard deviation (σ) across four repeated scans was calculated for 3D-PULSAR and IR-3D-PULSAR. In addition, the various scans were repeated on two volunteers on two different days to assess reproducibility.
All scans were performed on a Philips 3T Achieva scanner equipped with QUASAR dual mode gradients capable of maximum gradient amplitude of 80 mT/m and a slew rate of 200 T/m/s. A six-channel receive-only head coil was used for acquisition. Scan parameters were: TR/TD = 2380/1800 msec; 60 pairs of control/label images; tagging region width = 200 mm applied 20 mm inferior to imaging slab; data acquisition: 3D-TFEPI with 24 slices (31 slice encodings), fat suppression, 4 mm slice thickness, 80 × 80 matrix, SENSE factor = 2.5 along phase-encoding direction; DAC window ≈600 msec; bolus duration τ = 900 msec; scan time ≈5 min. The excitation pulses were Gaussian filtered sinc of duration 1.14 msec. Fat suppression was achieved by applying a pulse of duration 7.5 msec at off-resonance. Fat suppression pulse was employed once prior to acquisition of each 3D dataset. For IR-3D-PULSAR, the TI time was fixed at 920 msec prior to data acquisition. The manufacturer-supplied CLEAR option was used to correct for receive field inhomogeneity for all images.
A total of 12 healthy volunteers were scanned under an Institutional Review Board (IRB)-approved protocol. Informed consent was obtained from all volunteers after the nature of the procedure had been fully explained. Of the 12 volunteers, five were used for the background suppression study and the repeatability study; a different set of five volunteers was used only for the blurring reduction study; one was used only for the repeatability study only, while one was used for repeatability across different days along with another subject from the first group.
Table 1 shows the values for 〈σ(ΔM(t))〉/〈ΔM(t)〉 averaged over all five volunteers. Note that ΔM(t) value in CSF is small, resulting in higher ratios of 〈σ(ΔM(t))〉/〈ΔM(t)〉. A considerable reduction in this measure for GM, WM, and whole-brain perfusion signal with IR-3D-PULSAR as compared to 3D-PULSAR is seen, indicating better temporal stability. In addition, mean signal within CSF space was just 0.29 ± 0.76 with the background suppression inversion pulse as opposed to –0.93 ± 1.76, indicating better suppression of spurious perfusion signal in CSF.
|〈σ(ΔMGM)〉/〈ΔMGM〉||1.53 ± 0.87||1.13 ± 0.43|
|〈σ(ΔMWM)〉/〈ΔMWM〉||4.4 ± 2.72||2.43 ± 1.5|
|〈σ(ΔMCSF)〉/〈ΔMCSF〉||−7.12 ± 9.28*||11.67 ± 40*|
|〈σ(ΔMGlo)〉/〈ΔMGlo〉||0.31 ± 0.03||0.14 ± 0.04|
Figure 3 shows CBF maps (in mL/100g/min) for every fourth slice of the 24 slices acquired with 3D-PULSAR and the corresponding images with IR-3D-PULSAR. Figure 4 shows a plot of the GM CBF values obtained with the two sequences. The average GM CBF value across the five volunteers was 58.2 mL/100g/min with 3D-PULSAR and 58.9 mL/100g/min with IR-3D-PULSAR. Note that good correspondence between the two values indicates that the assumed value for inversion efficiency (η = 0.91) is accurate.
Figure 5 shows six transverse CBF image slices (out of 24) and reformatted sagittal and coronal slices obtained using the two different scans: constant flip angle (Fig. 5a,c) and modulated flip angle (Fig. 5b,d). As can be seen, CBF images obtained with the modulated excitation angle train appear sharper and suffer from lower partial volume effects. This is also reflected in the GM CBF values reported in Table 2. The mean value for GM CBF was 55.8 ± 8.4 and 65.9 ± 5.7 mL/100 g/min for the five volunteers before and after correction for blurring. There was a corresponding reduction in mean WM CBF values from 10.1 ± 2.4 to 8.0 ± 1.8 mL/100g/min indicating reduced partial volume effects after correction. A paired t-test between the CBF values shows significant differences (P < 0.05) for both GM and WM before and after blurring correction. The number of voxels classified as GM was also lower for the modulated flip angle case by an average of 20%, indicating reduced blurring.
|Volunteer||Constant flip||Variable flip|
|GM CBF||WM CBF||GM CBF||WM CBF|
Table 3 shows mean and standard deviation (σ) values across repeated scans for the globally averaged GM segmented images for the two cases: 3D-PULSAR and IR-3D-PULSAR. Figure 6 shows three select segmented GM CBF image slices from the 3D stack (row A) obtained with IR-3D-PULSAR; standard deviation images across four repeated scans for 3D-PULSAR (B) and (C) standard deviation images for IR-3D-PULSAR. (Window/level for images in (B) and (C) is the same but different from images in (A).) Images obtained with 〈M0A〉 for 3D-PULSAR show a similar qualitative appearance as images in Fig. 6b, indicating that the increased noise is not related to use of measured M0As vs. 〈M0A〉. Average σ across six volunteers is 3.93 when using individually measured M0As but is 3.15 with use of 〈M0A〉 for 3D-PULSAR, while it is just 2.51 for IR-3D-PULSAR. Average GM CBF values were within 3% for two volunteers across two different scanning sessions.
|1||57 ± 3.4||59 ± 3.9|
|2||53 ± 5.6||54 ± 0.9|
|3||64 ± 2.4||62 ± 1.3|
|4||58 ± 3.0||61 ± 2.2|
|5||59 ± 4.9||56 ± 2.6|
|6||71 ± 4.3||73 ± 4.2|
Our objective was to perform whole-brain perfusion imaging in a clinically reasonable scan time of ≈5 minutes. Pulsed arterial tagging along with 3D turbo field echo planar acquisition allows for sufficient SNR to achieve this goal. Although the data presented in this work were from 24 slice acquisitions, we have used the sequence provided here to acquire up to 35 slices of 4-mm thickness with an in-plane resolution of 2.9 × 2.9 mm in other ongoing studies. Use of an inversion pulse reduced temporal fluctuations in GM and WM and suppressed spurious signal from the CSF. It also showed improvement in CBF maps in areas of high susceptibility, as seen in Fig. 3. A shorter TR is possible even with a background suppression pulse, since a saturation pulse is applied to the imaging slab at the beginning of the spin labeling sequence. Extended data acquisition in the nonsteady-state can result in substantial blurring and result in reduced values for GM CBF as well as misclassification of GM CBF. Using a modulated flip angle train to account for this blurring gives higher values in GM by reducing partial volume effects.
The acquisition scheme used here (3D TFEPI) has advantages and disadvantages compared to 3D GRASE or spiral acquisition schemes. While single-shot spiral scanning (3D FSE stack of spirals) (6) can achieve a similar spatial coverage and resolution as the 3D TFEPI scheme, perfusion images will show increased off-resonance-related blurring. Spiral imaging is sensitive to gradient group delays, which can result in artifacts particularly when scanning at oblique angles. Compared to 3D TFEPI, 3D GRASE can offer improved off-resonance-response. However, since 3D TFEPI uses low flip angle excitation pulses and 3D GRASE depends on large flip angle excitation and refocusing pulses, the sampling efficiency in 3D TFEPI can be higher than that of 3D GRASE when selective pulses are used. This will to some extent improve the spatial resolution and/or coverage along the slice-encoding direction. However, based on a different pulse design (for example, by using nonselective refocusing pulses), it may be possible to make a 3D GRASE acquisition faster than a 3D TFEPI acquisition. By keeping echo times relatively short (TE ≈9 msec) with our TFEPI acquisition, the susceptibility effects are considerably reduced.
Values for bolus length and transit time were fixed based on empirical observations of the perfusion signal to allow sufficient SNR as well as to fulfill assumptions related to the general kinetic model. A bolus cutoff time of 900 msec ensures that even the fastest spins in the external carotid arteries (≈100 cm/s) have not yet completely left the tagging slab. Uncertainty in transit times can cause errors in the perfusion values. Using longer delay times can reduce the magnitude of these errors, at least for GM CBF. We eschewed the use of vascular signal crushers since the reduction in perfusion signal was considerable and the use of a sufficiently long delay time (1.8 seconds) did not warrant it. In patients with severely compromised flow, this will lead to increased vascular signal despite the prolonged delay time. Areas with prolonged arrival times will show reduced CBF values. WM perfusion values are generally unreliable with pulsed arterial sequences due to longer transit times. This violates the GKM model condition that requires that the entire bolus reach the imaging voxel under consideration. In addition, partial volume effects result in additional errors (21). Another recent study (22) has shown that about 45% of WM pixels show no statistically significant perfusion signal in a clinically feasible scan time of 5 minutes. Consistent with other studies dealing with pulsed arterial labeling, we found increased variability and decreased values for WM CBF (21, 23). As a result, except to show reduction in blurring, WM CBF values should not be taken as accurate.
A possible explanation of higher WM values reported in some other studies using pulsed labeling with similar or smaller bolus time could be contamination of perfusion signal from adjoining higher perfusion GM areas. Reducing the duration of the bolus while keeping delay time the same would result in an increase in the allowed bolus transit time and improve WM CBF values. In addition, reducing the bolus duration may reduce errors resulting from dispersion of the bolus. Reducing bolus duration, however, would impact the SNR of the perfusion signal and entail prolonged scan times. Increased bolus duration can also result in reduced perfusion values due to dispersion of the bolus. Typically, the effect is much less with the QUIPSS II-like sequence used here (24). Use of multiple inversion times (25) or a Look–Locker acquisition along with a model-free approach (10) can address the issue of different arrival times. Such an approach would compromise SNR, resolution, and spatial coverage and would therefore be more useful for CBF quantification in a few slices with 2D acquisition. Patients with severely compromised perfusion may still not be accurately quantified by the above methods.
We showed that as long as partial volume and saturation effects are avoided, M0A measurement does not significantly alter results when compared with other sources of error. Standard deviation maps for 3D-PULSAR reflect errors due to measurement of M0A as well as errors due to relatively poorer background suppression. M0A measurement errors result in a standard deviation that is higher than when M0A variation is absent (as when 〈M0A〉 is used with 3D-PULSAR). The average coefficient of variation (COV) defined as (σ/μ) is higher by about 19.8%. The fact that the σ map for 3D-PULSAR is more heterogeneous when compared with the IR-3D-PULSAR maps, particularly adjacent to CSF spaces, reflects the superior background suppression of IR-PULSAR; COV is further reduced by about 21% with IR-3D-PULSAR over 3D-PULSAR using 〈M0A〉. However, performing the Student's t-test reveals no significant difference between values obtained using individual M0A or 〈M0A〉 (t-value = 1.14 < 2.23 for a P-value of 0.05) for 3D-PULSAR, indicating that M0A measurements do not change results significantly.
GM CBF values were obtained by segmenting out regions of high perfusion from those that exhibit lower perfusion. This has some limitations and can result in cross-contamination of signal from WM and GM. A GM map could be independently derived from a higher-resolution scan and applied to the CBF map to determine perfusion signal originating in the GM after registration and correction for any motion between the two scans.
Despite the use of CLEAR to correct for receive field inhomogeneity, shading effects were seen in certain acquired datasets, a result of uncorrected transmit and receive field inhomogeneity. Correction for residual RF field inhomogeneity should mitigate the shading effect and result in improved perfusion images.
In conclusion, modifications to the PULSAR sequence permit whole-brain CBF measurements to be performed in an efficient manner. Further improvements in accuracy and efficiency are possible and are areas of future exploration. For example, multitransmit body coils can reduce errors introduced by B1 inhomogeneity. Improvements in coil design and parallel imaging algorithms can further reduce acquisition and echo time with subsequent reduction in EPI-related susceptibility effects. Scanning at higher fields (7T) can compensate for shorter acquisition time related SNR reduction.
The authors thank Dr. Xavier Golay, Esben Petersen, Maarten Versluis, and Dr. Frank Hoogenraad for initial help provided for the project.