Jessica L. Wisnowski, Rafael C. Ceschin and Vincent J. Schmithorst

Underlying Principles of DW-MRI Acquisition

At its core, DW-MRI involves the application of two additional pulses of magnetic field gradient (called “diffusion-encoding” or “diffusion-sensitizing” gradients) to a T2-weighted sequence after the excitation pulse but before the readout. During the first gradient pulse, spin precession is accelerated in accordance with the spatial position of the individual water molecules; spins associated with water molecules with a high Z coordinate, for example, will precess more quickly after administration of a gradient pulse along the Z direction, whereas spins associated with water molecules with a low Z coordinate will precess more slowly. Therefore the net effect of the first gradient pulse on the ensemble of spins is that the spins begin precessing at different rates and consequently “dephase,” resulting in signal attenuation. The second gradient pulse is equal in direction, magnitude, and duration (δ) to the first and is either of opposite polarity (in the case of a gradient echo acquisition) or of the same polarity (in the case of a spin echo acquisition) but placed after a 180-degree refocusing pulse. Assuming that the spins do not move from their original locations, the effect of the second gradient pulse will be to precisely undo the effect of the first and thus “rephase” the spins so there is no longer any signal attenuation due to spin dephasing. However, under physiologic conditions, water molecules possess thermal energy and therefore will move a finite distance away from their original locations during the time between gradient pulses (Δ). Thus the rephasing is incomplete and the signal will be attenuated compared with the signal with no diffusion-sensitizing gradients.

Assuming that the movement of water molecules is not hindered by any form of barrier (so-called “free diffusion”), the mean squared distance that the spins will move over a given period is described by the Einstein-Smoluchowsky equation and is linearly proportional to the time (Δ) and to the self-diffusion coefficient D. The amount of attenuation is a function of the gradient strength G, gradient duration δ, time between gradient pulses Δ, and diffusion coefficient D. Typically G, δ, and Δ are combined to derive the “b-value”; the higher the b-value, the greater the signal attenuation in the resultant DW images (Fig. 26-1). In fact, signal attenuation is exponential, where S is the measured signal and S₀ is the signal in the absence of diffusion weighting. As a result, D (measured in mm²/s) can be estimated by obtaining DW images at different b-values or by obtaining images with and without diffusion weighting.

However, the diffusion of water molecules in the brain differs in two important respects from free diffusion. First, the diffusion of water is hindered by a variety of barriers, including axon sheaths and glial cell and astrocyte membranes. Hence the measured diffusion coefficient is not a self-diffusion coefficient and therefore is referred to as an apparent diffusion coefficient (ADC). Furthermore, the diffusion of water in the brain is not isotropic (i.e., independent of direction). For instance, diffusivity along an axon direction will be larger than diffusivity perpendicular to the axon direction. Hence DW images often are obtained using a variety of gradient directions to infer information about the diffusivity of water molecules in different directions; the amount of attenuation in the DW images is dependent on the diffusion of water molecules only in the direction of the applied diffusion-encoding gradients.

It also is important to note that because DW measurements reflect an attenuation of signal at a given spatial location, maintenance of sufficient signal to noise in the resultant data is an inherent challenge in DWI. Moreover, the time needed to acquire data sufficient for some of the most advanced postprocessing techniques, which require acquisitions at many different gradient direction and b-values, is often well outside of what is typically feasible in clinical settings and in many pediatric populations. Thus in practice, most pediatric DW-MRI studies generally are limited to the more basic DW-MRI models (i.e., DTI, described in a later section of this chapter), although recent developments provide hope that other DW-MRI techniques soon will be clinically feasible.

Diffusion Weighting, Apparent Diffusivity, and Their Application in Clinical Settings

Clinically relevant information is available even from a single DWI acquisition. For instance, diffusion is restricted in regions of cytotoxic edema after a stroke, and these regions therefore will be hyperintense on DW images. However, typical DWI acquisitions are strongly T2-weighted, because a long echo time is necessary as a result of the time needed to apply the diffusion-sensitizing gradients. Therefore it is important to distinguish hyperintensity on DWI that represents true diffusion restriction from hyperintensity that reflects tissue T2 prolongation (often termed “T2 shine-through”). This differentiation usually is performed by the additional acquisition of an image without diffusion weighting (e.g., b = 0) and quantifying ADC on a pixel-by-pixel basis, which subsequently can be represented in gray scale as an ADC map. Additionally, instead of a single DW acquisition, three DW images typically are acquired using three orthogonal gradient directions, and the results are averaged to obtain directionally averaged DW images and ADC maps to minimize the effects of anisotropy. The directionally averaged ADC map is proportional to the trace of the diffusion tensor (described later), and thus the DWI images and ADC maps often are referred to as “trace-weighted DWI” and “trace diffusion tensor maps,” respectively (Fig. 26-2).

As previously noted, in the past two decades, trace DW images with corresponding ADC values have demonstrated remarkable clinical utility in the detection of acute cerebral ischemia, often before such injuries otherwise become apparent. In persons with ischemia, a critical drop in cerebral blood perfusion leads to energy failure, and more specifically, a failure of the Na+/K+–adenosine triphosphatase pumps in the cell membrane. This phenomenon, in turn, leads to an influx of sodium (and other ions) and water into the cell, causing the cell to swell (i.e., cytotoxic edema). Although other events also might contribute to the change in ADC, it has been suggested that ADC is most sensitive to a small change in the distribution of water between extracellular and intracellular environments, and thus ADC can be viewed as a marker of fluid-electrolyte homeostasis.

In addition to fluid homeostasis and intracellular water, DWI and ADC also are sensitive to the relative properties of water in the extracellular space. Thus DWI also is useful in the differential diagnosis of intracranial lesions. In lesions with high cellularity (e.g., high-grade tumors), the increased cellularity restricts water motion in the extracellular space, resulting in decreased ADC in the lesion (or in areas of the lesion with relatively higher cellularity). Accordingly, DWI often can distinguish high-grade tumors such as primitive neuroectodermal tumors from other lower grade pediatric brain tumors such as ependymomas or astrocytomas (Fig. 26-3).

Diffusion Tensor Imaging

Although three diffusion-sensitizing gradients are sufficient for calculating a directionally averaged ADC image, a minimum of six directions is needed to characterize diffusion anisotropy. Anisotropic diffusion is directionally dependent, as within white matter fiber bundles, and is distinguishable from isotropic diffusion, which is observed in free fluids (e.g., the lateral ventricles) (Fig. 26-4).