Exogenous tracer methods for perfusion MRI use a model assuming that the tracer is confined to the intra-vascular compartment and does not diffuse to the ex-travascular space [1, 3-5].
Imaging can be obtained either dynamically (rapid imaging over time after a bolus injection) or, less frequently, in the steady state (after constant infusion has achieved an equilibrium concentration of the tracer in the blood) [1-5].
With steady-state techniques, a baseline image is acquired before injection and a post-infusion image is obtained up to 30 min after slow administration of the agent (i.e. during the steady state) . Subtraction of the baseline from the post-contrast image allows a map of absolute blood volume to be obtained. This method is not very common due to its inherently low signal-to-noise ratio (SNR) and the fact that patient movement between scans can affect accuracy .
Dynamic susceptibility contrast perfusion MRI (DSC), also called dynamic contrast-enhanced perfusion MRI, dynamic susceptibility-weighted perfusion imaging, perfusion-weighted imaging, or first-pass bolus tracking perfusion MRI, is the most widespread, clinically applicable MR technique for estimating cerebral haemodynamic parameters, especially in stroke, but also for tumour imaging and other research applications.
This technique is based on the passage of a bolus of contrast medium through the arterial and capillary circulation and on the transient changes it produces in vessels and surrounding tissues (Fig. 9.1).
Under normal perfusion conditions, in the presence of an intact blood brain barrier (BBB), the contrast medium remains confined within the vascular network and does not diffuse in the extravascular space. While passing through the cerebral vasculature, a short bolus of contrast material produces local magnetic field inhomo-geneities that lead to a reduction in the transverse relaxation time of the tissue. This susceptibility effect can be recorded by a series of ultra-fast T2*- or T2-weighted sequences using gradient-echo and spin-echo sequences, respectively. The signal intensity-time curves can be converted to concentration-time curves, which allow the calculation of haemodynamic parameters such as blood volume, blood flow and transit time (Fig. 9.2).
In pathological perfusion conditions, e.g. hyper-acute ischaemic lesion, the signal reduction is attenuated or delayed whereas in brain tumour the signal varies in relation to grade.
Perfusion imaging requires: (1) rapid administration of the contrast agent; (2) acquisition of ultra-fast sequences; and (3) post-processing of the native images to calculate the intensity-time curve and obtain perfusion maps.
1. Infusion of the gadolinium-chelated agent must be as rapid as possible to ensure correct mixing of the agent with blood and a sharp signal drop profile. This requires using a large intravenous line (18 or, preferably, 16 gauge), an automatic injector and high infusion velocity (5 ml/s) and doses of
2. Imaging sequences must be sufficiently fast to allow accurate measurement of the rapidly changing signal from the first pass of the bolus and have adequate temporal resolution (<2 s for the entire brain) . The most widely applied imaging sequences for these studies are T2*-weighted gradient-echo single-shot echo-planar imaging (EPI) and T2-weighted spin-echo EPI.
Spin-echo EPI sequences display excellent sensitivity for the susceptibility effect produced by 5 ^m capillaries. Gradient-echo EPI sequences are generally more sensitive up to a vessel diameter of 7 ^m. Spin-echo sequences are thus inherently weighted towards the microvasculature and are therefore less sensitive to larger vessels. Gradient-echo EPI sequences are currently those used most frequently [37-40]. They are characterized by superior sensitivity in detecting the signal change produced by the passage of the contrast material and therefore require smaller doses of the
agent. Their drawback is their high sensitivity not only to the microvasculature but also to the macro-vasculature, including large arteries and veins on the brain surface. Susceptibility effects are elicited not only by the passage of contrast material, but also by haemoglobin degradation products, the interfaces between different tissues and age-related iron accumulation in the extrapyramidal nuclei. These effects are much more pronounced at high field intensities .
An alternative to T2- or T2*-sequences are T1-weighted 3D spoiled gradient-echo sequences, which give rise to relaxivity, rather than susceptibility effects. The relaxivity effect of paramagnetic contrast agents results in a shortening of T1 relaxation time (yielding higher signal on T1), whereas the susceptibility effect shortens both T2 and T2* (giving lower signal on T2 or T2*) .
Because the relaxivity effects are much stronger than the susceptibility effects, T1-weighted sequences require a smaller amount of contrast agent (approximately 10%) compared with T2 or T2* weighted sequences, allowing multiple repeated studies . The main disadvantage of T1-weighted techniques is that the effects of BBB disruption are much greater than with T2- and T2*-weighted sequences [1, 3].
3. The decrease in signal intensity produced by the passage ofabolus of gadolinium chelate allows a signal intensity-time curve to be obtained by calculating the change in signal intensity in a single voxel (or in a single region of interest) as a function of time. This curve is then converted to an agent concentration-time curve.
The sum of all the concentration-time curves of all the voxels in a given slice generates perfusion maps from which various haemodynamic parameters can be calculated, including cerebral blood volume (CBV), cerebral blood flow (CBF), mean transit time (MTT) and time-to-peak (TTP) (Figs. 9.3,9.4). These parameters are dependent on the specific features of the bolus injection (injection rate, amount and concentration of contrast material) and on specific variables of the patient being imaged (total body vascular volume and cardiac output) [1,3]. As a result, haemodynamic parameters cannot be directly compared between different subjects and may even differ between examinations of the same individual performed at different times [1, 3].
Thus, relative values can be obtained using an internal standard of reference such as normal-appearing greyorwhite matter (Figs. 9.5,9.6) [1,3]. For diffuse processes, in which the internal reference may also be affected, absolute quantitation is required. Absolute quantitation of CBV and CBF has been attempted using methods that measure arterial input to the brain, but their accuracy is debated [1, 3].
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