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Currently, most MR perfusion studies are acquired on 1.5 T machines, but MR systems operating at higher field strengths are increasingly becoming available in clinical practice [4, 5, 8]. In general, imaging at higher magnetic field strengths offers at least a linear SNR increase, but its utilization is impaired by problems related to magnetic susceptibility artefacts.

Since magnetic susceptibility increases with field strength, at 3.0 T image distortion maybe critical, particularly if EPI pulse sequences are used, as in most protocols for DSC perfusion studies. In fact, with EPI sequences the trade-off is increased B0 inhomogeneity and magnetic susceptibility differences at air-tissue interfaces, leading to signal drop-out and geometric distortions, most notably in the phase encode direction [53].

Image quality may thus be severely impaired by distortion and blurring around tissues interfaces, especially in patients who have undergone neurosurgery, and assessment of signal intensity at the passage of a bolus of contrast agent maybe inaccurate.

Other drawbacks inherent with T2* susceptibility are systematic overestimation of perfusion because of the different relaxivity of the paramagnetic contrast agent in the larger vessels (from which the input function is obtained) and in capillaries; problems in defining a good input function because of low spatial resolution on T2*-weighted EPI; and finally problems in assessing a damaged BBB. The latter is a critical element in the evaluation of brain tumour perfusion, because relaxivity is different when the contrast agent is confined to a small compartment as opposed to distributed in the interstitial space [54].

In order to minimize these negative effects, use of rapid 3D fast field-echo T1-weighted or 3D FLASH T1-weighted sequences has recently been proposed [54, 55]. Although acquisition is slower and yields less spa tial and temporal resolution than EPI, FLASH seems to suffer less from spatial distortion than EPI, thus yielding a better arterial input function [55].

On the other hand, T2*-weighted contrast-enhanced perfusion imaging using 3.0 T systems offers some potential advantages. In particular, shorter T2 and T2* relaxation times and increased SNR [56, 57] may allow a greater T2* signal drop to be obtained for a given amount of contrast bolus during capillary passage [58].

A recent study of the feasibility of MR perfusion at 3.0 T demonstrated that image quality of the perfusion source images was not impaired, leading to rating distortion and blurring as minor in most images [8].

This was also true of critical anatomical areas, such as the posterior fossa and the regions close to the base of the skull (e.g. hippocampus and brainstem) [8]. This is probably due to use of a multi-shot echo-shifted three-dimensional gradient-echo echo-planar imaging sequence, in which the short echo train length may compensate for the field-dependent increase in susceptibility effects [8, 59-62].

Choice of correct echo time may also affect the performance of DSC perfusion imaging. A recent work comparing different echo times (ranging from 21 to 45 ms) in DSC perfusion in 17 patients demonstrated that the shortest echo time used yielded the best images [63].

Important benefits have recently been reported from using EPI sequences in DSC perfusion at 3.0 T with parallel imaging [37], also with the implementation of a multi-channel coil array [53], or with use of spin echo EPI, as an alternative to gradient echo EPI [64].

Parallel imaging (PI) techniques such as simultaneous acquisition of spatial harmonics (SMASH), sensitivity encoding (SENSE) or generalized autocalibrat-ing partially parallel acquisition (GRAPPA) allow the shortening of scanning time by reducing the number of phase-encoding steps needed. Correct image reconstruction is achieved by inclusion of additional spatial information obtained from the spatial variation of coil sensitivity of multiple receiver coils. The reduction in encoding time can be used to achieve greater temporal resolution or to increase spatial resolution of the otherwise typically low-resolution dynamic scans [37].

The use of EPI sequences at 3.0 T is ideally suited for combination with PI techniques because the reduced echo train length results in several improvements.

The first benefit is reduction of image distortion and blurring from B0 inhomogeneities produced by the interfaces between tissues and tissue boundaries or by haemorrhage products and surgical material after cra-niotomy. These image distortions are proportional to the off-resonance frequency produced by the inhomo-geneities and inversely proportional to the time necessary to traverse the fc-space. With PI the sensitivity of

EPI to any of these off-resonance artefacts, particularly significant at 3.0 T, can be reduced [37].

Another advantage is the reduction of the influence of T2* relaxation on spatial resolution, also referred to as fc-space filtering. fc-space filtering limits the resolution that can be achieved due to signal loss during spatial encoding. At 1.5 T a typical contrast agent reduces the T2* time of brain tissue to 20-40 ms. Therefore encoding steps performed after the actual T2* time contribute little to the fc-space information and the image appears as if acquired at much lower resolution. Because of the shorter T2* relaxation at higher field inten sities, this effect is even more pronounced at 3.0 T. This effect is also very important for the measurement of the arterial input function. With conventional EPI, the signal inside vessels vanishes completely and the arterial input function can be detected only around vessels.

Moreover, reduced shot duration in PI allows the acquisition of more slices in the same scanning time, enabling increased coverage of the region being investigated.

Finally, using interleaved EPI sequences, application of PI is also recommended for selecting an appropriate echo time.

Fig. 9.8. DSC at 3.0 T. Comparison between two different concentrations ofthe same dose of contrast agent: 0.5 mmol (a, c) versus 1.0 mmol (b, d). At 3.0 T the enhanced magnetic susceptibility requires lower doses and/or lower concentrations than 1.5 T systems. (Courtesy of General Electric)

The dose of contrast agent also needs to be adjusted to the high-field strength setting. Although structural contrast-enhanced brain imaging at 1.5 T is usually performed with 0.1 mmol/kg of body weight, 0.2 mmol of gadolinium chelate is the most widely used dose and is considered optimal for DSC perfusion studies [65 - 71], also in case of higher concentrations of gadolinium-based contrast agents (e.g. gadobutrol, Gado-vist; Schering, Berlin, Germany) [8, 72]. When this dose (0.2 mmol) was used for DSC perfusion at 3.0 T, the stronger susceptibility effects caused a complete signal void during the first pass of gadolinium, especially in grey matter, affecting the accuracy of signal drop calculation and impairing sensitivity to perfusion deficits [8]; for this reason, a dose of 0.2 mmol is not recommended for high-field perfusion imaging.

A recent 3.0 T study comparing different doses (0.05, 0.1 and 0.2 mmol) has recommended a dose of 0.1 mmol on the basis both of subjective analysis of the quality of the perfusion maps and of a quantitative assessment of perfusion variables [8] (Fig. 9.8).

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