High-Resolution Imaging of the Brain 18

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R. Bammer, PhD; S. Nagle, MD, PhD

Department of Radiology, Stanford University, Lucas Center, 1201 Welch Road, Stanford, CA 94305-5488, USA

C O N T E N T S

18.1 Introduction 183 18.2 Structural MRI 185 18.3 MR Angiography 190

18.4 Quantitative/Functional MRI 193 18.5 Pediatric MRI 196

18.6 Conclusion 196

References 197

High-Resolution Imaging of the Brain 18

Roland Bammer and Scott Nagle

18.1

Introduction

Resolution enhancement in MRI is of great potential for increased diagnostic accuracy in brain and spine imaging. With the advent of high-ﬁ eld systems increased signal-to-noise ratio (SNR) affords smaller voxel sizes, but overall scan time is still a limiting factor for high-resolution brain imaging in a clinical setting. Parallel imaging is of great beneﬁ t since it allows us to achieve high-resolution 2D and 3D acquisitions in clinically acceptable time frames. In addition, parallel imaging can also diminish the amount of image blurring and geometric distortions leading to obvious resolution and quality improvements without altering the acquisition matrix size.

This chapter critically addresses the general advantages and limitations of high-resolution neuroimaging in concert with the additional capacity provided by parallel imaging. Speciﬁ cally, the role of parallel imaging in high-resolution structural MRI, magnetic resonance angiography, and functional MRI in the broader sense (i.e., diffusion and perfusion MRI as well as classical functional MRI) are discussed.

The improved scanning effi ciency of parallel imaging methods can be applied in a number of fruitful ways in the fi eld of brain MR imaging. As described in the fi rst part this book, these parallel-imaging strategies cleverly incorporate the spatially varying sensitivity profi les of multiple-channel receive coils in order to reduce the number of k-space measurements necessary to reconstruct an image (Hutchin- son and Raff 1988; Kwiat et al. 1991; Sodickson and Manning 1977; Pruessmann et al. 1999). In conventional Cartesian k-space sampling schemes, this is typically realized by decreasing the number of phase- encoded steps. In cases of high SNR, the reduction of phase-encoded steps by a factor R (accompanied by its obligatory R decrease in SNR Pruessman et al. 1999) can be used to increase either spatial resolution or temporal resolution. These advantages are not mutually exclusive and it depends on the specifi c scan protocol whether one favors more rapid scanning or higher resolution. In addition, the artefact and blurring associated with several multiple-echo or long-readout sequences can be mitigated by parallel-imaging techniques (cf. Chap. 10).

Increasing the spatial resolution may allow the detection of smaller lesions, may better characterize the internal structure of larger lesions (e.g., calciﬁ - cation, blood products, demyelination, cystic components, etc.), and may better delineate the lesion boundaries with respect to normal anatomy, improving the accurate localization of a lesion (especially important in the prepontine, suprasellar, cerebellopontine angle, cavernous sinus, orbital, and Meckel’s cave regions). The imaging of white matter disease, stroke, neoplasm, and vascular disease could all ben-

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efi t from these advantages. Pediatric and neonatal brain imaging also demands fast, high-resolution imaging because of the relatively small brain size and the diffi culties in keeping a child still throughout the scan. Similar considerations apply also for imaging the spinal cord. Three-dimensional spoiled gradient-echo sequences, used in the evaluation of mesial temporal sclerosis in the work-up of seizures, tumor treatment planning, and voxel-based morphometry in neurodegenerative disorders, could benefi t from the use of parallel imaging in both phase-encoded directions to further increase resolution without increasing scan time (Weiger et al. 2002a).

Conversely, shortened scan times alone can increase patient throughput, resulting in obvious operational and patient comfort improvements. Simply decreasing scan time reduces the risk of patient motion degrading a study. A number of other creative methods for further reducing motion artefacts through the use of parallel imaging have been proposed and demonstrated (Bammer et al. 2004; Kuhara and Ishihara 2000;

Bydder et al. 2002, 2003; Atkinson et al. 2004).

Parallel imaging also yields advantages in other, less obvious ways. Conventional spin-echo T2- weighted imaging requires prohibitively long scan times due to the long TRs needed, resulting in the widespread adoption of fast spin-echo (FSE) or turbo spin echo (TSE) techniques. In order to increase their efﬁ ciency T2-weighted FSEs are usually carried out with relatively long echo trains. Unfortunately, such multi-echo sequences suffer from blurring and decreased SNR relative to conventional spin-echo sequences, due to the T2-decay occurring while the multiple echoes are being acquired (cf. Chap. 9).

By using parallel-imaging methods to decrease the

required number of phase-encoded measurements, the effective TE can be reduced by decreasing the echo-train length, simultaneously generating both improved SNR and decreased image blurring due to the fact that data in the periphery of k-space are being acquired earlier under the T2-decay curve (Fig. 18.1). Alternatively, for a given echo-train length the FSE scans can be made more effi cient and fewer TR intervals are required to complete the acquisition of the required k-space information. Even there, the faster traversal through k-space can lead to reduced T2-blurring. Similarly, by reducing the read-out time, off-resonance artefacts can be reduced (Bammer et al. 2001a). In pulse sequences which are SAR-limited, such as FLAIR or single-shot FSE scans, parallel imaging offers the advantage of eliminating a large fraction of the RF excitation pulses and is particularly attractive at higher fi eld strengths when other SAR-diminishing techniques, such as hyperechoes (Hennig and Scheffl er 2001) or variable-rate selective excitation (VERSE, Conolly et al. 1988), are not an option.

The application of parallel imaging to intracranial MR angiography (MRA) offers many advantages and is discussed in detail in Chap. 26. Wilson et al. (2004) also provide an excellent review of the use of parallel imaging in MRA. Due to the small size of the vessels, and the presence of intravoxel dephasing on time- of-ﬂ ight (TOF) images, high resolution is desired.

Phase-contrast (PC) techniques must obtain fourfold more data than TOF if full 3D velocity measures are desired, leading to long scan times, especially if full brain coverage at high resolution is desired. Fortu- nately, MRA images (especially contrast-enhanced MRA images) often have a high baseline SNR and

Fig. 18.1a–c. Example of k-space ﬁ lter effect due to T2-decay in FSE. a During normal FSE readout the net magnetiza- tion decays with T2 (T2* in EPI) which in turn leads to a T2-dependent weighting of k-space that ultimately results in blurred images. b Using parallel imaging (R=2 in this example), the echo-train length can be shortened by 1/R. c As every other line in k-space (for the R=2 case) is skipped, the relevant k-space information is obtained much faster and the signal loss towards the edge of k-space is reduced (indicated by the dotted line). While the acquisition matrix remains the same, the apparent resolution is improved by blurring reduction.

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could beneﬁ t from the application of parallel imaging to increase spatial resolution at the expense of a slight SNR loss, especially when moving to 3 T with its approximately doubled SNR (Weiger et al. 2000, 2002a) relative to 1.5 T.

Quantitative MR imaging methods, such as diffusion-weighted (DWI), diffusion tensor (DTI), perfusion-weighted imaging (PWI), and functional MRI (fMRI), rely for the most part on single-shot echo planar (EPI) pulse sequences (Mansﬁ eld 1977). The EPI pulse sequences offer especially attractive targets for the application of parallel imaging for much the same reasons as the multi-echo sequences described above. Due to their inherently long read-out times, EPI methods are prone to T2*-related signal decay and an extreme sensitivity to susceptibility effects (cf.

Chap. 10). These effects manifest in blurred and geo- metrically distorted images (especially in regions of the brain near the skull base and paranasal sinuses).

Often these regions of the brain are precisely the areas of interest in the clinical settings of stroke or when using fMRI to map cortical brain activity. In the setting of DTI for the purposes of white matter tractogra- phy, e.g., high spatial resolution, is critical for resolving regions where fi ber tracks cross and to minimize partial-volume effects and geometric distortions of the tracts. Without parallel imaging it is diffi cult to acquire high-resolution DTI data in a reasonable amount of time. Fortunately, with EPI techniques the inherent R signal loss of using parallel imaging is mitigated by the fact that the signal is acquired earlier under the T2*-decay curve (at a shorter effective TE), resulting in a signal increase (cf. Chap. 10). The benefi ts of using parallel imaging in EPI applications have been suc- cessfully demonstrated in various studies (Bammer et al. 2001a, 2002; Griswold et al. 1999; Jaermann et al.

2004; Kuhl et al. 2005).

18.2

Structural MRI

Increasing the spatial resolution of MR scans, both in-plane and through-plane, offers the opportunity to better visualize structural changes and characterize individual lesions. Here, especially, MRI’s diagnostic sensitivity to small lesions can be increased, and the conspicuity of small structural details in larger lesions (such as small calciﬁ cations or lipid deposits) may increase the diagnostic speciﬁ city. Although the

exact role of higher-resolution in neuroradiology, especially in the context of varying levels of experience of individual readers, is thus far not known, it is reasonable to expect that the degree of conﬁ dence and the speed of interpretation increase with increasing spatial resolution.

In white matter diseases, such as multiple sclerosis (MS), appreciation of the internal complexity of these lesions can help to differentiate stages of inﬂ ammation, demyelination, and gliosis in the individual plaques, which is important in clinical management and the evaluation of response to therapy (Fig. 18.2).

A recent study has shown that high-resolution structural MRI in MS patients promises to be more sensitive than conventional MRI (Erskine et al. 2005).

This may be due to lesions being obfuscated by small vessels, conﬂ uent signal abnormalities, or partial- volume averaging. In some cases MS plaques can also be found in cortical and juxtacortical regions (Kidd et al. 1999; Kangarlu et al. 2004). Especially for cortical lesions, the lesion conspicuity is often impaired by different relaxation rates and signiﬁ cant partial- volume contamination; thus, high-resolution imaging can add to a higher diagnostic certainty.

Improved spatial resolution is also helpful for tumor imaging in regions of the brain where precise spatial localization is important: the cerebellopontine angle; the prepontine cistern; the suprasellar region;

the cavernous sinus; the foramen of the ventricular system (Monroe, Magendie, and Lushke); and Meckel’s Cave, in particular. The fi rst step towards characteriz- ing an intracranial lesion is the accurate localization into intraaxial or extraaxial compartments as the differential diagnoses for lesions in these compartments are distinct. These compartments are separated by the dura, and the higher the resolution, the easier it is to directly visualize the dural membranes and accurately localize the lesion. Even the differentiation of the most typical extraaxial lesions, such as men- ingioma and schwannoma, can be facilitated by an increased conspicuity of their diagnostic hallmarks (e.g., dural tail, calcifi cation, lipid deposits). Further- more, the internal structure of a neoplastic lesion has important implications in terms of differential diagnosis and is often best appreciated on higher-resolution images, with their greater anatomic detail and reduced partial-volume averaging. The presence of small fl ow voids, calcifi cations, lipid deposits, cystic lesions, and overall texture of a lesion may only be detectable at higher resolution and offer the possi- bility of increasing the specifi city of the differential diagnosis reported by the radiologist (Fig. 18.3).

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In cases in which the diagnosis is known and the purpose of the study is for preoperative planning (e.g.,

“cyber knife” stereotactic radiosurgery), precise defi nition of the lesion boundary is quite important in order to avoid damage to critical nearby structures. Even if the spatial resolution of most radiosurgery methods is beyond what conventional MRI provides, a clear demarcation between neoplastic tissue and adjacent non-neoplastic tissue is of great benefi t both for treatment planning and for post-treatment follow-up for recurrent disease. For diffusely infi ltrating tumors, such as glioblastoma multiforme, a higher spatial resolution might help differentiate between edematous and neoplastic tissue. Extracranially, the accurate diagnosis and preoperative planning for head and neck tumors relies on very accurate identifi cation of the tissue planes trans- gressed by the tumor and would benefi t from improved high-resolution imaging (cf. Chap. 19).

From an applications perspective, parallel imaging allows scanning with conventional resolution in less time or scanning with increased resolution without increasing scan time. These approaches are

not mutually exclusive, as parallel imaging is best described as a tool to most efﬁ ciently acquire images using phased-array coils that are used for comple- mentary image encoding. Due to the sequential image formation process in MRI, and the nature of Fourier encoding itself, the increase of spatial resolution is accompanied by a proportional increase in the number of phase-encoded steps. This in turn can increase the overall scan time quite dramatically, which is often forgotten in the excitement to move toward higher resolution. Aside from the scan time penalty, quadrupling the SNR loss when doubling the in-plane resolution must be affordable by the protocol, i.e., possible within reasonable scan times and supported by a considerable baseline SNR to begin with. Signal averaging can be used to compensate the SNR penalty to some extent but one has to keep in mind that, for example, compensating for a 4× loss in SNR it would require 16 averages to maintain equal SNR and thus excessively prolongs scan time.

Increasing resolution along the frequency-encoding direction (at a ﬁ xed gradient strength and band-

Fig. 18.2a–d. High-resolution (0.43×0.43- mm² in-plane resolution, 3 mm slice thickness) MRI at 1.5 T in a patient suffering from a relapsing/remitting form of multiple sclerosis with an enhanc- ing lesion in the right periventricular area. a,b T2-weighted FSE scans. c,d Proton-density-weighted FSE scans. The improved appreciation of the internal complexity of these lesions might help to differentiate stages of inﬂ ammation, demyelination, and gliosis in the individual plaques. Enlarged perivenular spaces and associated focal lesions can be seen adjacent to the posterior horns of the lateral ventricles.

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width or water/fat shift per pixel) is accompanied by a proportional increase of the echo readout duration, which in turn prolongs the echo-train length of all FSE and EPI-based methods. For several sequences the consequences of larger echo spacings and overall echo train lengths are a reduced number of slices that can be interleaved per TR interval and a more signiﬁ cant k-space ﬁ ltering effect due to the more pronounced T2-decay (Fig. 18.1). If possible, extended echo spacings are often remedied by increasing the readout gradient and bandwidth at the expense of a reduced SNR; however, with some methods the maximum gradient strength and slew rate are already at their limits (e.g., balanced-SSFP, EPI) and only fractional echoes or parallel imaging is a viable alternative. Increasing the resolution along the phase-encoding direction is very costly in terms of additional scan time.

With the introduction of new clinical 3 T systems and their higher baseline SNR, the use of larger acquisition matrices is becoming a viable option again. As the ﬁ eld strength increases from 1.5 to 3 T, the effec-

tive SNR gain doubles to what would otherwise be achieved only by a tedious averaging process at 1.5 T.

As a rule of thumb one has to scan four times longer at 1.5 than at 3 T to achieve comparable SNR levels.

Of course, relaxation and B₁ effects will mitigate this beneﬁ t slightly. Firstly, the prolonged tissue relaxation times require longer TR to provide adequate signal recovery. Secondly, B₁ inhomogeneities can lead to incomplete or partial excitation and refocusing as well as reduced receive sensitivity. Particularly for high ﬁ eld applications, parallel imaging is a very attractive option since it helps to keep high-resolution MRI within clinically acceptable time frames, and despite its inherent SNR loss still provides diag- nostically adequate scans (Fig. 18.4).

The development of new RF coil designs has followed the introduction of parallel imaging (cf.

Chap. 14). Speciﬁ cally, multi-channel array coils, having between 8 and 16 independent receiver coils distributed around the subjects head, have been developed for brain imaging. In research studies,

Fig. 18.3a–d. A 3D T1- weighted spoiled gradient-echo sequence (TR/

TE=12/4.5 ms; D=20q;

0.86×0.86×1.5 mm³) at 3 T with 2D-GRAPPA R=2u2 along both phase- encoding directions.

Follow-up exam of a patient suffering from an anaplastic astrocy- toma after partial resec- tion, radiotherapy, and chemotherapy (arrow).

High-resolution structural scanning was accomplished almost four times (plus time to acquire auto-calibra- tion lines) faster than with conventional image encoding. a Mid-sagit- tal pre-contrast view. b Sagittal pre-contrast cut through the tumor. c,d Adjacent axial slices after contrast administration show edema (asterisk) and residual contrast enhancement (arrow).

(Courtesy R. Stollberger and F. Payer, Medical University Graz, Austria)

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scans have even been performed with prototype arrays containing more than 100 independent coils.

In general, these array coils provide better SNR than comparable birdcage coils and boost regionally the SNR by factors up to two (Bammer et al. 2001b). At our institution we also made the observation that at 3 T the signal pile-up due the B₁-fi eld focusing effect is less pronounced with these phased-array coils than with standard birdcage coils. Most likely, the higher coil sensitivity at the brain’s periphery (i.e., these areas are closer to the coils) offsets the fi eld focusing effect. Parallel imaging can be implemented in several ways to improve high-resolution structural brain imaging. The most important aspect is, of course, the reduction of overall scan time when other imaging options, such as partial k-space acquisition, fractional (i.e., rectangular) fi eld of view (FOV), or reducing the number of signal acquisitions are already exhausted or are not an option. Given enough baseline SNR, parallel imaging allows one to acquire high-resolution scans within reasonable acquisition periods. Fur-

thermore, parallel imaging’s impact on image deblur- ring can also be considerable even without changing the acquisition matrix size. Here, the faster traversal through k-space leaves the signal less time to decay with T2 or T2*; thus, in addition to more rapid image formation, images will appear to have improved resolution because of the diminished blurring effects.

An easily overlooked problem related to increasing in-plane resolution (especially in slice-selective MRI scans) is the frequent anisotropic voxel size. To pre- vent this voxel anisotropy from becoming even more exaggerated as in-plane resolution is increased it is paramount to reduce the slice thickness accordingly;

however, this often reduces the available SNR even further and also requires an increased number of slices to cover a speciﬁ c area.

For applications with a large number of slices and short repetition times (and when isotropic voxels and diminished partial-volume averaging are important) the use of 3D acquisitions with phase-encoding along two principal axes is advisable (see Fig. 18.3). Three-

Fig. 18.4a–f High-resolution imaging of the brain and intracranial vessel at 1.5 T (top row) and 3.0 T (bottom row) in a normal volunteer using identical acquisition parameters. Comparison of GRAPPA-induced noise enhancement using reduction factors of R=1 (a,d), R=2 (b,e), and R=3 (c,f). Parallel imaging clearly beneﬁ ts from the higher baseline SNR and the spatially more distinct RF characteristics at 3.0 T. The SNR measurements revealed comparable values for R=3 at 3.0 T and R=1 at 1.5 T.

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dimensional imaging is frequently used for volumet- ric studies, stereotactic therapy planning, and angiographic studies. Typically, T1-weighted 3D spoiled gradient-echo sequences or MPRAGE sequences are used for image acquisition. Classically, 3D MRI is a very attractive method as it has increasing SNR ben- eﬁ ts over 2D imaging with an increasing number of slices to cover the brain. While 2D acquisitions receive signal only from a thin slice, the signal in 3D acquisitions emanates from the entire slab.

In addition to the higher baseline SNR, 3D MRI provides the opportunity to perform parallel imaging along both phase-encoding directions independently.

The overall scan acceleration will be the product of the accelerations achieved in each of the two phase-encoding directions. For the same overall acceleration factor it has been shown that acceleration in both dimen- sions provides a more benign g-factor-related noise enhancement than in only one dimension (cf. Chap. 3).

This can be easily understood if one recalls the basic principle of SENSE. With increasing reduction factor, R, the distance between voxels that are aliased on top of each other will decrease. A reduced distance between aliased voxels also implies a less distinct coil sensitivity variation which makes it harder for the reconstruction to separate these voxels. An overall scan time reduction by a factor three or four can be achieved very easily with parallel imaging along both phase-encoded directions and has enormous consequences for sequences that run otherwise 15 min or more. In research settings, reduction factors of 8 and higher have been reported for 3D acquisitions; and some researchers presented results where the reduction factor even exceeded the

theoretical limits of R (i.e., the number of independent receiver coils) by capitalizing on special mathematical regularization methods (Fig. 18.5; Katscher 2003).

Although there is currently no strict convention in place, it is important to report the acceleration factor in both phase-encoded directions separately, i.e., R_3D=3×2 for a twofold acceleration in the in-plane and a three- fold acceleration in the through-plane phase-encoded direction. It is noteworthy that residual aliasing along the phase-encoding direction perpendicular to the slice is much harder to detect than the more obvious artefacts seen from in-plane reconstruction errors. If in doubt, a reformation of the data to an orthogonal plane might reveal the problem.

As already mentioned, imaging at higher resolution affords a better characterization of morphological abnormalities. Particular areas in which high resolution is of potential beneﬁ t is the hippocampus area, the pituitary region, the orbits, or the cranial nerves;

however, the small FOVs that are generally used for these studies might pose a fundamental limit for parallel imaging and the currently available coils. In this context it is important to understand that a high degree of scan acceleration is only feasible if there is a signiﬁ cant variation in coil sensitivity across the FOV.

If the spatial coil sensitivity variation becomes spatially less distinctive, it becomes more difﬁ cult to cor- rectly separate the aliased voxels. The individual coil sizes as well as their size and orientation relative to the prescribed FOV are essential parameters and characterize the capacity of the parallel-imaging method to remove aliasing. This information is also matched by the spatially varying g-factor. In other words, with

Fig. 18.5a–c. Example for a SENSE reconstruction with two surface coils and a reduction factor R=3. a Example image serving as input for the simulation (TSE/turbo factor 30, TE=100 ms, TR=2000 ms, pixel volume 0.9×0.9×8 mm³, D = 60°, B₀=1.5T).

b Reconstruction result for R=3 using a uniform k-space density. c Reconstruction result for the same parameters as for b, but using a Gaussian k-space density. (Courtesy of Dr. Katscher, Philips Research Labs, Hamburg, Germany)

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diminishing FOVs and without adaptation of the individual coil element size, parallel imaging may no longer keep up in its ability to unfold aliased voxels.

This might eventually become problematic for areas in the middle of the brain remote from all coil elements and at high acceleration factors.

18.3

MR Angiography

Angiographic sequences, in particular contrast- enhanced MRA, are generally not limited by SNR, and are therefore ideally suited for combining them with parallel imaging. As with conventional structural MRI, the ability to further reduce acquisition time can be beneﬁ cial not only in reducing patient motion but also in providing larger FOV or better spatial and temporal resolution; the latter enables the use of time-resolved MRA acquisitions that provide important functional information in addition to morphological details, rep- resenting a major advantage of MRI over advanced multi-detector CT angiography. Increased in-plane anatomic coverage in intracranial MRA afforded by parallel imaging allows one to display a larger part of the circle-of-Willis (COW), including ACA, MCA, and PCA and the distal portion of both the basilar artery

and the ICAs. Increasing the coverage in cranio-caudal direction makes it also easier to prescribe the posi- tion of the image stack and to more accurately cover the targeted vessels. A shorter imaging time should also facilitate the use of this technique in restless and uncooperative patients, as in the setting of acute stroke or when imaging children.

Time-of-fl ight (TOF) MRA is based on fl ow effects and is one of the primary diagnostic tools for patients with suspected intracranial aneurysms and steno-occlusive disease. In particular, gradient-echo sequences are used for TOF MRAs, in which the infl ow of fresh, unsaturated blood into RF-saturated tissue leads to increased signal intensity within vessels. Due to the small caliber of intracranial vessels and the presence of vessels throughout the brain, high-resolution imaging over larger FOVs is a prerequisite for an accurate diagnostic work-up. Specifi cally, high spatial defi - nition is relevant for the imaging of both steno-occlusive disease and aneurysms where the visualization of fi ne anatomic details in the COW, such as small vessel branches and the neck of aneurysms, is needed. Like- wise, the high spatial defi nition is advantageous in the characterization of AVMs (Fig.18.6). In this context, reducing the voxel size also diminishes the inherent sensitivity to intravoxel dephasing in TOF and further improves the quality of TOF MRA.

As with structural MRI, MRA also beneﬁ ts from the SNR gain afforded by the migration to higher

Fig. 18.6. Sagittal maximum intensity projection of a high-resolution axial T1-weighted 3D gradi- ent-echo TOF MRA (acquisition matrix 832×1024; R=2.5) at 3 T in a patient suffering from an arte- riovenous malformation (arrowheads) in the right insula region and a superﬁ cial venous drainage into the superior sagittal sinus (arrow).Increased baseline SNR at 3 T affords high-quality MRAs in this patient with great conspicuity of the ateriovenous malformation. The acquisition of 150 slices took 5:12 min. (Courtesy W. Willinek, Universiyt of Bonn, Bonn, Germany)

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magnetic ﬁ eld strengths. In addition, the saturation effect of the surrounding parenchyma, and hence increased vessel-to-parenchyma CNR, is even more pronounced at 3 T due to the prolonged T1 relaxation times of semisolid tissues relative to essentially no change in T1 of blood. The extra SNR can be invested in high spatial resolution protocols, improving the depiction of small vessel segments (Willinek et al.

2003, 2004; Bernstein et al. 2001; Al-Kwiﬁ et al.

2002; Thomas et al. 2002), and, in turn, improving image quality of intracranial aneurysms (Metens et al. 2000) and the diagnostic accuracy of detecting cerebrovascular disease. Since typical imaging times for high-resolution protocols are overly long, parallel imaging is ideally suited to reduce imaging time.

A recent study demonstrated the ability of parallel imaging to both reduce acquisition time and to improve the anatomic coverage for high spatial resolution TOF MRA at 3.0 T with comparable image quality. In this particular study, acquisition times were reduced from 7:57 to 5:12 min using reduction factors of 2.5, while the anatomic coverage was increased 1.5 times. In a series of 80 acute stroke patients we also observed that the overall quality of TOF MRA as well as the visualization of distal branches of the MCA, ACA, and PCA improved substantially after beginning to use multi-element head coils. This can most likely be attributed to the increased peripheral receive sensitivity of such coils.

Phase-contrast (PC) MRA (Dumolin et al. 1989) is another angiographic MR method and utilizes motion-probing gradients to provide fl ow-related contrast along the direction of the fl ow-encoding gradient. The PC MRA is important for evaluating lesions were blood deposition products appear hyperintense in TOF-MRAs (and therefore mimic infl ow) or for abnormalities with slow fl ow where the TOF mechanism is lost by saturation effects (i.e., where not enough fresh blood replenishes saturated blood to provide suffi cient vessel-to-parenchyma contrast).

In addition, with PC-MRA, the velocity-encoding parameter can be used to better discriminate between fast-fl owing arteries and the slow fl ow of blood in the veins. One application in which PC MRA is frequently used is MR venography. Unfortunately, PC MRA is a notoriously slow imaging technique. This is because at least two or up to four measurements are required to measure fl ow along one direction or three directions (plus a reference measurement). Parallel imaging can substantially speed up PC MRA. Here, the abundant SNR inherent in PCA supports the use high reduction factors. Furthermore, in 3D PC MRA, parallel imag-

ing can be applied in both the phase and slice-encoding directions, allowing even higher reduction factors.

Recently, it has been shown that reduction factors of up to 68 are possible in PC MRA of the brain with isotropic 1-mm³ resolution. Without parallel-imaging scan-time reduction, such an examination would take about 40 min to complete, whereas with parallel imaging it is possible to reduce the overall acquisition time to 57 min (Fig. 18.7).

At our institution we have recently begun to perform 3D PC MRA to measure all three components of ﬂ ow at several instances within the RR interval(R.

Bammer et al., submitted). Such measurements provide us detailed information of blood ﬂ ow within a 3D volume, temporally resolved across the cardiac cycle.

The knowledge of magnitude and orientation of ﬂ ow in space and its change over time throughout the RR interval allows us to provide streamline analysis and vir- tual particle tracing in intracranial vessels (Fig. 18.8).

Due to this complex, multi-dimensional acquisition scheme, it is understandable that acquisition times are considerable. By means of GRAPPA we could cut down the acquisition time substantially. At 3 T a reduction factor of 3 was feasible without noticeable reconstruction artefacts or noise enhancement and without the use of contrast material (see Fig. 18.4).

Fig. 18.7. Axial projection MIP of high-resolution whole-brain 3D MR venography at 3 T using phase-contrast MRA and SENSE in two directions. A SENSE reduction factor of 4 (2×2) was taken in slice and phase-encoding direction to acquire a true 512 matrix within 1:40 min. The high spatial resolution MRA allows for depiction of the ophthalmic arteries (arrowheads).

(Courtesy of W. Willinek, Universiyt of Bonn, Bonn, Germany)

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The use of contrast material has clearly revolu- tionized the role of MRI in angiography. To perform contrast-enhanced (CE) MRA, a 3D spoiled gradient-echo sequence covering the anatomy of interest is synchronized to the arrival of a T1-shorten- ing contrast-agent bolus. The use of CE MRA for neurovascular studies requires relatively short and exactly timed acquisitions to provide maximum arterial enhancement and avoid venous overlay. Cap- turing an arterial phase can either be accomplished by rapid acquisition of temporal dynamics (thus including at least one frame of pure arterial phase) or precisely timing a higher-resolution scan to the

arrival of the bolus (using “ﬂ uoroscopic triggering”

or other timing mechanisms; Fig. 18.9). A high temporal resolution is beneﬁ cial to CE MRA because it provides dynamic visualization of contrast kinetics that can include important information for diagnosis of certain vascular pathologies. Certainly, there is a trade-off between achievable temporal and spatial resolutions but parallel imaging can be used to improve temporal resolution without compromis- ing resolution or artefact level. Recently, Golay et al.

(2001) showed the utility of parallel imaging in time- resolved CE MRA of carotid arteries. Using R=2 they acquired a 270×65×50-mm³ volume at a resolution

Fig. 18.8. Time-resolved (20 time points/RR interval) high-resolution streamline visualization of the left internal carotid artery, the middle cerebral artery, and the anterior cerebral artery. The time-resolved 3D velocity data were computed from high-resolution 3D phase-contrast MR scans with fl ow measurement along the principal axes. Streamline computation was performed by fourth-order Runge-Kutta numerical integration of the fl ow vector fi eld. The streamlines are color coded [ranging from 50 cm/s (red) to 0 cm/s (blue)] based on the magnitude velocity. The images are arranged from top to bottom beginning with peak systole and ending with late diastole. The pulse wave traveling distally can be clearly appreciated.

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of 1.0×1.3×2.4 mm³ in under 6 s. They found that this temporal resolution was sufﬁ cient to provide at least one dynamic acquisition with pure arterial phase without the need for timing the acquisition to peak arterial signal. Further improvements may be anticipated in this area, especially from combining parallel imaging with non-Cartesian trajectories, such as TRICKS (Turski et al. 2001) or VIPR (Du et al. 2004).

18.4

Quantitative/Functional MRI

Both quantitative and functional MRI methods usually require rapid imaging methods, such as EPI (Mansﬁ eld 1977) or spiral imaging (Meyer et al. 1992). Unfortunately, image quality and spatial resolution with EPI and spiral imaging is usually frustrated by geometric distortions, signal loss, and image blurring. These artefacts emanate from signal sources with off-resonant spin precession rates (rela-

tive to that of water), and the amount of phase offset that these spins accrue during the relatively long EPI and spiral readouts (cf. Chap. 10). A signiﬁ cant magnetic ﬁ eld inhomogeneity across a voxel can lead to incoherent phase accrual and leads to intravoxel dephasing and signal loss. Local susceptibility gradients, such as those around the auditory canals or the frontal sinuses, can also lead to signal pile-up artefacts. In EPI, these artefacts occur along the phase- encoding direction, whereas in other techniques, such as spiral or radial imaging, they can manifest as radial blurring. Despite the fact that EPI is an ultra- fast MRI technique, the sampling of data along the phase-encoded direction is still slow (on the order of 1 ms/phase-encoded step) and gives rise to the afore- mentioned artefacts (Farzaneh et al. 1990). Signal T2* decay limits the possible achievable readout time and hence the ultimate resolution.

A strategy to reduce these susceptibility artefacts as well as those related to image blurring is to traverse k-space as fast as possible. In other words, the higher the k-space velocity (i.e., 'k/'t), the lower are the artefacts. Interleaved trajectories in EPI and spiral imaging have been used to speed up the k- space traversal; however, the time to form an interleaved echo-planar image increases by the number of interleaves used times the respective repetition time. This fact limits the use of multi-shot methods for DWI and PWI since they normally require single- shot methods. Speciﬁ cally, PWI methods that rely on the measurement of the bolus passage through the brain require high temporal resolution in order to follow contrast agent passage reliably through the brain. EPI is currently the only method that is capable of providing adequate temporal resolution.

In DWI, random bulk physiological motion during the diffusion-encoding phase leads to unpredictable phase accrual that may vary from shot to shot and is compounded with the intentional phase encoding.

Using single-shot EPI guarantees an equal motion- induced phase error for each phase-encoding step that will vanish after magnitude calculation. Other than that, DWI theoretically does not require rapid image formation (if the phase perturbation can be corrected), although it is our experience that the likelihood of patient motion and blurring increases with prolonged scan times. Also, new advances in DTI require large numbers of individual DWI scans (Liu et al. 2004) and would lead to prohibitively long scan times if not carried out with fast imaging methods.

With parallel imaging, artefacts in EPI and spiral imaging can be reduced quite signiﬁ cantly. When

Fig. 18.9. Sagittal MIP of high-resolution coronal 3D contrast- enhanced MR angiography of the supra-aortic arteries using SENSE in two directions. A SENSE reduction factor of 4 (2×2) was taken in slice and phase-encoding direction to acquire 365 slices (matrix: 512×512 before zero-ﬁ lling). Note that the entire supra-aortic arteries are displayed, including cervical, facial, and occipital branches (arrows). (Courtesy W. Willinek, University of Bonn, Bonn, Germany)

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we ﬁ rst set out to use parallel imaging for our diffusion-weighted EPI methods (Bammer et al. 2001a) we aimed to accelerate k-space traversal–similar to interleaved EPI (Butts et al. 1994) acquisitions – to increase the bandwidth per pixel and hence diminish artefacts while maintaining a single-shot regime to provide high temporal resolution and avoid the need for phase navigation (Anderson and Gore 1994). The theoretical SNR loss in these sequences is mitigated by the shorter EPI readout, which leaves more time for the diffusion encoding (Bammer et al. 2001a). This in turn leads to shorter echo times and less T2 signal decay. In a standard Stejskal-Tanner DWI sequence, the effective time available for diffusion-encoding gradients to be played out is determined by the time between the 180q refocusing pulse and the begin of the EPI readout.

Since the TE is determined by the time the center of k-space is traversed, the beginning of the EPI readout will signiﬁ cantly protrude into the second diffusion- encoding interval leaving only a signiﬁ cantly shortened interval open for diffusion encoding.

To improve the efﬁ cacy of DWI partial k-space acquisition (i.e., projection onto convex sets (POCS) or homodyne reconstruction) is normally used to reduce the number of gradient echoes required before the formation of the spin echo. The efﬁ cacy can be

Fig. 18.10a,b. Diffusion-weighted imaging of a patient suffering from multiple ischemic lesions in different vascular territories. a Image quality on conventional single-shot EPI is frustrated by blurring and susceptibility artefacts. b Signiﬁ cantly better image quality is achieved by adding parallel imaging. In this case (R=3), the shortened readout afforded to increase the acquisition matrix from 128 to 192 with still reduced levels of geometric distortions. Small lesion conspicuity was considerably better due to the increased matrix size and the diminished T2*-related blurring.

further improved by the use of parallel imaging since it essentially reduces the EPI train by the factor R.

In addition, the shortened EPI readouts allow one to interleave more slices per TR or to reduce overall acquisition time, since often the EPI readout over all slices determines the minimum TR.

Using a GRAPPA-like approach at 1.5 T we were able to use reduction factors up to four to reduce distortions and simultaneously almost double the standard acquisition matrix to provide DWI scans at signiﬁ cantly improved spatial resolution (Fig. 18.10). Preliminary results from an NIH-funded study in 50 consecutive stroke patients showed that when SENSE and conventional single-shot EPI-based DWI were compared by rating the quality of scans from 1 (technically inad- equate) to 5 (perfectly diagnostic scan). The mean score for conventional scans was 3.5 (r0.7), whereas the score for SENSE DWI was only slightly higher with 3.8 (r1.0); however, on a “winner-take-all” basis the SENSE DWI outperformed the conventional scan in 54%, performed equivalently in 16%, and underper- formed in 30% of the cases. These results are slightly less optimistic than the study published recently by Kuhl et al. (2005) but still support the important trend towards improved image quality and diagnostic capacity with parallel imaging-enhanced DWI.

b a

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In this context, we emphasize that SENSE DWI will change DWI’s diagnostic sensitivity and speciﬁ - city. The DWI is already a very sensitive methodol- ogy; however, the parallel-imaging-related resolution enhancement and the minimization of susceptibility artefacts will increase the conﬁ dence level of individual readers. Given the fact that most of the predi- lection sites for susceptibility distortions are known, it is rather unlikely to misclassify such hyperintense areas as stroke, although we have to admit that certain signal hyperintensities in the posterior fossa and the temporal lobe are sometimes quite challenging and require cross-referencing with other series and experience from other DWI scans to rule out ischemia.

We instead expect that the hyperintensities from susceptibility distortions might mask truly existing ischemic changes and that these confounders can be minimized by applying parallel imaging to DWI.

We also expect that the resolution enhancement will increase the diagnostic sensitivity to lacunar infarc- tion and small embolic lesions; the latter are of great diagnostic relevance if they can be detected among larger pre-existing ischemic lesions and even more so in a different vascular territory/hemisphere; however, further study is needed in a much larger cohort in order to achieve enough statistical power to investi- gate how frequently diagnosis and treatment would be altered based on SENSE DWIs. This study is currently underway at our institution. Preliminary results show that parallel-imaging-enhanced DWI provides higher diagnostic sensitivity and conﬁ dence for small lesions, but in several embolic cases multiple, larger

lesions have already provided enough evidence for the neuroradiologist to reach a diagnosis.

Both arterial spin labeling (ASL; Detre et al. 1994) and dynamic susceptibility contrast (DSC) (Rosen et al. 1989; Weiskoff and Rosen 1992) based perfusion- weighted MRI (PWI) will benefi t similarly from parallel imaging if the data acquisition is performed with EPI or spiral MRI (which is normally the case; Reishofer et al. 2002; Bammer and Moseley 2004). In a recent study we were able to apply reduction factors of four to DSC-PWI scans without any apparent reconstruction artefacts other than the obligatory parallel-imaging-related noise enhancement (Fig. 18.11). Due to the signifi cantly reduced EPI readout length, the overall brain coverage could be considerably increased. The improved resolution of the DSC-PWI scans led to an enhanced differentiation between cortical gray matter and juxta-cortical white matter in processed parameter maps [i.e., cerebral blood fl ow (CBF) or cerebral blood volume (CBV)]. Another artefact source in DSC- PWI that parallel imaging addresses very well is the magnetic susceptibility change that occurs during bolus passage. Because of the low bandwidth per pixel, the signifi cant susceptibility changes can shift signal sources in large arteries by more than a pixel which makes the measurement of the arterial input function (AIF) extremely diffi cult (Rausch et al. 2000). Here, the increased bandwidth per pixel due to parallel imaging helps to reduce these shifts and allows for a more accurate measurement of the AIF.

The effects and consequences of applying parallel imaging to functional MRI are very similar to those

Fig. 18.11a,b. Dynamic-susceptibility contrast enhanced perfusion-weighted MRI. a T2*-weighted single-shot EPI with R=4 (top) and R=3 (bottom). The ﬁ rst column shows one of 15 baseline scans. The second column shows the maximum contrast change during the bolus passage. Despite the high acceleration, both data sets are of high quality without noticeable reconstruction artefacts. b Cerebral blood ﬂ ow (CBF) maps calculated from R=3 (top) and 4 (bottom) scans using a deconvolution approach in combination with automatic AIF detection. The CBF maps obtained with R=4 are slightly noisier than R=3 maps.

a b

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of PWI (Golay et al. 2004). Within the past few years several studies have investigated the use of parallel imaging in fMRI, with reductions in distortion as their primary goal. Weiger et al. (2002b) implemented a spiral-SENSE technique and used it in visual, motor, and taste functional experiments. In their experiments, they found that the SNR and signal-to-ﬂ uctuation-noise ratio (SFNR) were generally diminished by 20 and 13%, respectively, when using a reduction factor R=2. It is important to note, however, that parallel imaging did not affect the detection power of the activation, and the number of activated voxels remained more or less con- stant over the experiments. Another important ﬁ nding of Weiger et al.’s (2002b) study was the ability of parallel imaging to recover part of the signal loss present in the deep orbito-frontal areas. This is a very critical area to evaluate in several functional tasks. Its ability to be reliably scanned with fMRI is of great value.

In a similar study, Preibisch et al. (2003) examined the detection power using simple motor tasks relative to the reduction factor. They found that at intermedi- ate reduction factors (R=2) they were able to increase the number of slices per unit of time at high spatial resolution, as well as decrease the distortions without signifi cant loss in statistical power. At higher reduction factors, they experienced a rapid loss in SNR and loss in statistical power. Schmidt et al. (2003) and Morgan et al.(2004) found optimal reduction factors of 23 for SENSE EPI at 3.0 T, whereas Little et al. (2004) found no detrimental use in activation detection with R=3 using GRAPPA. In this context, it is important to point out that in fMRI the temporal stability ultimately determines the statistical signifi cance of the fMRI activation. In general, this stability is determined by a sum of the physiological noise variance and image noise variance (Kruger and Glover 2001). In each parallel-imaging-enhanced scan the measured fMRI signal is certainly affected by the g-factor-related noise enhancement; however, if the physiological noise is the dominant noise source, especially with increasing fi eld strength, the application of parallel imaging and hence the altered image noise will not much infl uence the sensitivity of the fMRI experiment.

18.5

Pediatric MRI

In general, the same considerations that apply to high- resolution structural MRI on the adult side apply

also to the pediatric population; however, pediatric neuroimaging is further challenged by the smaller patient size as well as by the lack of cooperation that one can anticipate from these patients; the latter often requires the need for sedation or even anesthesia.

Pediatric parallel imaging has therefore important beneﬁ ts for pediatric patients. Capitalizing on the faster acquisition speed, patient studies can be performed much faster and children have to spend less time in the magnet. This will either shorten the sedation/anesthesia time or simply reduce patient anxiety.

Of course, the gain in speed can be also invested in increasing the spatial resolution or to reduce geometric distortions. Like in adults the improved resolution might provide better lesion conspicuity or reveal hallmarks of disease otherwise unseen or equivocal.

So far, very little has been published on pediatric MRI in combination with parallel imaging. One limitation might be the limited availability of pediatric parallel imaging coils. Due to the wide range of head sizes seen in children and adolescents it is often difﬁ cult to provide size adequate head arrays. Cur- rently, only a few vendors provide phased-array coils for pediatric imaging. A combined head and spine array coil is currently available for purchase from one coil manufacturer for only two size groups (group 1:

up to 10 kg; group 2: 10–22.5 kg). All other patients have to be imaged with standard adult equipment.

Compared with the adult hardware the pediatric head arrays have fewer coil elements, thus limiting the maximum reduction factor. Comprehensive neuroexams often also include spectroscopy, which requires special postprocessing for multi-coil acquisition. Unfortunately, some MR vendors do not fully support this function in their products and therefore some pediatric neuroradiologists are reluctant to use multi-element coils. Switching coils during the study might be an alternative but is more complicated in pediatric examinations because of the frequent need for anesthesia and the associated ramiﬁ cations.

18.6 Conclusion

Neuroimaging is often challenged by the lack of detail in conventional structural and functional MRI scans as well as in MR angiography. The complexity of the brain’s morphology and the need for early and accurate diagnosis motivates the push to higher resolution.

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The sequential order of the MR acquisition proce- dures and the underlying biophysical MR properties that characterize tissue are signiﬁ cant speed-limiting factors in this otherwise very powerful modality. The introduction of parallel imaging has had a dramatic impact on the MRI community. Aside from acceler- ated image acquisition, it has been responsible for a remarkable overhaul of RF receiver technology (coils, electronic), which might have otherwise happened but probably not to such an extent.

Neuroimaging beneﬁ ts from a deliberate and well- balanced application of parallel imaging in several ways. Firstly, parallel imaging provides enough speed beneﬁ ts to acquire larger image matrices within a reasonable time. Secondly, it helps to diminishing blurring and geometric distortions. Thirdly, parallel imaging can be of utility to reduce motion artefacts and therefore improve the apparent resolution of neuroimaging studies. The true value of parallel imaging to high-resolution brain imaging can only be estimated. No conclusive studies are currently available and it also depends on the preferences of an individual radiologist whether or not resolution enhancement is truly relevant for the majority of examinations. It is a complex interplay between eco- nomic considerations, artefacts, and SNR that radi- ologists and institutions will have to balance as this technology continues to evolve.

Acknowledgements. This work was supported in part by the NIH (1R01EB002771, 1R01EB002771S3), the Center of Advanced MR Technology at Stanford (P41RR09784), Lucas Foundation, and Oak Founda- tion.

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