Neuroimaging: Structural Imaging: Magnetic Resonance Imaging, Computed Tomography

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Chapter 33A Neuroimaging

Structural Imaging: Magnetic Resonance Imaging, Computed Tomography

Computed Tomography

Computed tomography (CT; other terms include computer assisted tomography [CAT]) has been commercially available since 1973. The term tomography (i.e., to slice or section) refers to a process for generating two-dimensional (2D) image slices of an examined organ of three dimensions (3D). CT imaging is based on the differential absorption of x-rays by various tissues. X-rays are electromagnetic waves with wavelengths falling in the range of 10 to 0.01 nanometers on the electromagnetic spectrum. X-rays can also be described as high-energy photons, with corresponding energies varying between 124 and 124,000 electron volts, respectively. X-rays in the higher range of energies, known as hard x-rays, are used in diagnostic imaging because of their ability to penetrate tissue yet (to an extent) also be absorbed or scattered differentially by various tissues, allowing for the generation of image contrast.

Owing to their high energy, x-rays are also a form of ionizing radiation, and the health risks associated with their use, although minimal, should always be accounted for in diagnostic imaging. The x-rays generated by the x-ray source of the CT scanner are shaped into an x-ray beam by a collimator, a rectangular opening in a lead shield. The beam penetrates the slab of tissues to be imaged, which will absorb/deflect it to a varying degree depending on their atomic composition, structure, and density (photoelectric effect and Compton-scattering). The remaining x-rays emerge from the imaged slab and are measured by detectors located opposite the collimator. In fourth-generation CT scanners, the detectors are in a fixed position and the x-ray source rotates about the patient. As the beam of x-rays is transmitted through the imaged body part, sweeping a 360-degree arc for each slice imaged, the emerging x-rays are collected, then a computer analyzes the output of the detectors and calculates the x-ray attenuation of each individual tissue volume (voxel).

The degree of x-ray absorption by the various tissues is expressed and displayed as shades of gray in the CT image. Darker shades correspond to less attenuation. The attenuation by each voxel of tissue is projected on the flat image of the scanned slice as a tiny quadrilateral, generally square, called a pixel or picture element. Depending on the reconstruction matrix, a slice will be represented by more or fewer pixels, corresponding to more or less resolution. The shade of gray in each pixel corresponds to a number on an arbitrary linear scale, expressed as Hounsfield units (HU). This number varies between approximately −1000 and 3000+, with values of greater magnitude corresponding to tissues or substances of greater radiodensity, which are depicted in lighter tones. The −1000 value is for air, 0 is for water. Bone is greater than several hundred units, but cranial bone can be 2000 or even more. Fresh blood (with a normal hematocrit) is about 80 units, fat is −50 to −80. Tissues or materials with higher degrees of x-ray absorption, shown in white or lighter shades of gray, are referred to as hyperdense, whereas those with lower x-ray absorption properties are hypodense; these are relative terms compared to other areas of any given image.

By changing the settings of the process of transforming the x-ray attenuation values to shades on the grayscale, it is possible to select which tissues to preferentially display in the image. This is referred to as windowing. Utilizing a bone window, for instance, is very useful for evaluating fractures in cases of craniofacial trauma (Fig. 33A.1).

In CT, imaging contrast agents are frequently used for the purpose of detecting abnormalities that cause disruption of the blood-brain barrier (BBB) (e.g., certain tumors, inflammation, etc.). The damaged BBB allows for the net diffusion of contrast material into the site of pathology, where it is detected; this is referred to as contrast enhancement. Contrast materials used in CT scanning contain iodine in an injectable water-soluble form. Iodine is a heavy atom; its inner electron shell absorbs x-rays through the process of photoelectric capture. Even a small amount of iodine effectively blocks the transmitted x-rays so they will not reach the detector. The high x-ray attenuation/absorption will result in hyperdense appearance in the image. Other CT techniques requiring contrast administration are CT angiography, CT myelography, and CT perfusion studies.

More than 20 years ago, a fast-imaging technique called spiral (or helical) CT scanning was introduced to clinical practice. With this technique, the x-ray tube in the gantry rotates continuously, but data acquisition is combined with continuous movement of the patient through the gantry. The circular rotating path of the x-rays, combined with the linear movement of the imaged body, results in a spiral or helix-shaped x-ray path, hence the name. These scanners can acquire data rapidly, and a large volume can be scanned in 20 to 60 seconds. This technique offers several advantages, including more rapid image acquisition. During the short scan time, patients can usually hold their breath, which reduces/minimizes motion artifacts. Timing of contrast bolus administration can be optimized, and less contrast material is sufficient. The short scan time, optimal contrast bolus timing, and better image quality are very useful in CT angiography, where cervical and intracranial blood vessels are visualized. These images can also be reformatted as 3D views of the vasculature, which are often displayed in color and can be depicted along with reformatted bone or other tissues in the region of interest (Fig. 33A.2).

Superfast CT scanners have become available in the past 5 years. By multiplying by 4 the number of detectors, they can obtain 64 slices of an organ in a fraction of a second. They are particularly useful in cardiology and also allow for the acquisition of perfusion images of the entire brain. One shortcoming is a greater exposure to ionizing radiation per scan.

Magnetic Resonance Imaging

Basic Principles

Magnetic resonance imaging (MRI) is based on the magnetic characteristics of the imaged tissue. It involves creation of tissue magnetization (which can then be manipulated in several ways) and detection of tissue magnetization as revealed by signal intensity. The various degrees of detected signal intensity provide the image of a given tissue.

In clinical practice, MRI uses the magnetic characteristics inherent to the protons of hydrogen nuclei in the tissue, mostly in the form of water but to a significant extent in fat as well. The protons spin about their own axes, which creates a magnetic dipole moment for each proton (Fig. 33A.3). In the absence of an external magnetic field, the axes of these dipoles are arranged randomly, and therefore, the vectors depicting the dipole moments cancel each other out, resulting in a zero net magnetization vector and a zero net magnetic field for the tissue.

This situation changes when the body is placed in the strong magnetic field of a scanner (see Fig. 33A.3, A). The magnetic field is generated by an electric current circulating in wire coils that surround the open bore of the scanner. Most MRI scanners used in clinical practice are superconducting magnets. Here the electrical coils are housed at near–absolute zero temperature, minimizing their resistance and allowing for the strong currents needed to generate the magnetic field without undue heating. The low temperature is achieved by cryogens (liquid nitrogen or helium). Most clinical scanners in commercial production today produce magnetic fields at strengths of 1.5 or 3.0 tesla (T).

When the patient is placed in the MRI scanner, the magnetic dipoles in the tissues line up relative to the external magnetic field. Some dipoles will point in the direction of the external field (“north”), some will point in the opposite direction (“south”), but the net magnetization vector of the dipoles (the sum of individual spins) will point in the direction of the external field (“north”), and this will be the tissue’s acquired net magnetization. At this point, a small proportion of the protons (and therefore the net magnetization vector of the tissue) is aligned along the external field (longitudinal magnetization), and the protons precess with a certain frequency. The term precession describes a proton spinning about its own axis and its simultaneous wobbling about the axis of the external field (see Fig. 33A.3, B). The frequency of precession is directly proportional to the strength of the applied external magnetic field.

As a next step in obtaining an image, a radiofrequency pulse is applied to the part of the body being imaged. This is an electromagnetic wave, and if its frequency matches the precession frequency of the protons, resonance occurs. Resonance is a very efficient way to give or receive energy. In this process, the protons receive the energy of the applied radiofrequency pulse. As a result, the protons flip, and the net magnetization vector of the tissue ceases transiently to be aligned with that of the external field but flips into another plane, thereby transverse magnetization is produced. One example of this is the 90-degree radiofrequency pulse that flips the entire net magnetization vector by 90 degrees to the transverse (horizontal) plane (Fig. 33A.4). What we detect in MRI is this transverse magnetization, and its degree will determine the signal intensity. Through the process of electromagnetic induction, rotating transverse magnetization in the tissue induces electrical currents in receiver coils, thus accomplishing signal detection. Several cycles of excitation pulses by the scanner with detection of the resulting electromagnetic signal from the imaged subject are repeated per imaged slice. This occurs while varying two additional magnetic field gradients along the x and y axes for each cycle. Varying the magnetic field gradient along these two additional axes, known as phase and frequency encoding, is necessary to obtain sufficient information to decode the spatial coordinates of the signal emitted by each tissue voxel. This is accomplished using a mathematical algorithm known as a Fourier transform. The final image is produced by applying a gray scale to the intensity values calculated by the Fourier transform for each voxel within the imaging plane, corresponding to the signal intensity of individual tissue elements.

T1 and T2 Relaxation Times

During the process of resonance, the applied 90-degree radiofrequency pulse flips the net magnetization vectors of the imaged tissues to the transverse (horizontal) plane by transmitting electromagnetic energy to the protons. The radiofrequency pulse is brief, and after it is turned off the magnitude of the net magnetization vector starts to decrease along the transverse or horizontal plane and return (“recover or relax”) toward its original position, in which it is aligned parallel to the external magnetic field. The relaxation process, therefore, changes the magnitude and orientation of the tissue’s net magnetization vector. There is a decrease along the horizontal or transverse plane and an increase (recovery) along the longitudinal or vertical plane (Fig. 33A.5).

To understand the meaning of T1 and T2 relaxation times, the decrease in the magnitude of the horizontal component of the net magnetization vector and its simultaneous increase in magnitude along the vertical plane should be analyzed independently. These processes are in fact independent and occur at two different rates, T2 relaxation always occurring more rapidly than T1 relaxation (Fig. 33A.6). The T1 relaxation time refers to the time required by protons within a given tissue to recover 63% of their original net magnetization vector along the vertical or longitudinal plane immediately after completion of the 90-degree radiofrequency pulse. As an example, a T1 time of 2 seconds means that 2 seconds after the 90-degree pulse is turned off, the given tissue’s net magnetization vector has recovered 63% of its original magnitude along the vertical (longitudinal) plane. Different tissues may have quite different T1 time values (T1 recovery or relaxation times). T1 relaxation is also known as spin-lattice relaxation.


Fig. 33A.6 This diagram illustrates the simultaneous recovery of longitudinal magnetization (T1 relaxation) and decay of horizontal magnetization (T2 relaxation) after the RF pulse is turned off.

(Reprinted with permission from Hashemi, R.H., Bradley, W.G., Lasanti, C.J., 2004. MRI—The Basics. 2nd ed. Lippincott Williams & Wilkins.)

While T1 relaxation relates to the longitudinal plane, T2 relaxation refers to the decrease of the transverse or horizontal magnetization vector. When the 90-degree pulse is applied, the entire net magnetization vector is flipped in the horizontal or transverse plane. When the pulse is turned off, the transverse magnetization vector starts to decrease. The T2 relaxation time is the time it takes for the tissue to lose 63% of its original transverse or horizontal magnetization. As an example, a T2 time of 200 ms means that 200 ms after the 90-degree pulse has been turned off, the tissue will have lost 63% of its transverse or horizontal magnetization. The decrease of the net magnetization vector in the horizontal plane is due to dephasing of the individual proton spins as they precess at slightly different rates owing to local inhomogeneities of the magnetic field. This dephasing of the individual proton magnetic dipole vectors causes a decrease of the transverse component of the net magnetization vector and loss of signal. T2 relaxation is also known as spin-spin relaxation. Just like the T1 values, the T2 time values of different tissues may also be quite different. Tissue abnormalities may alter a given tissue’s T1 and T2 time values, ultimately resulting in the signal changes seen on the patient’s MR images.

Repetition Time and Time to Echo

As mentioned before, the amount of the signal detected by the receiver coils depends on the magnitude of the net magnetization vector along the transverse or horizontal plane. Using certain operator-dependent parameters, it is possible to influence how much net magnetization strength (in other words, vector length) will be present in the transverse plane for the imaged tissues at the time of signal acquisition. During the imaging process, the initial 90-degree pulse flips the entire vertical or longitudinal magnetization vector into the horizontal plane. When this initial pulse is turned off, recovery along the longitudinal plane begins (T1 relaxation). Subsequent application of a second radiofrequency pulse at a given time after the first pulse will flip the net magnetization vector that recovered so far along the longitudinal plane back to the transverse plane. As a result, we can measure the magnitude of the net longitudinal magnetization that had recovered within each voxel at the time of application of the second pulse, provided that signal acquisition is begun immediately afterwards. The time between these radiofrequency pulses is referred to as repetition time, or TR (Fig. 33A.7). It is important to realize that contrary to the T1 and T2 times, which are properties of the given tissue, the repetition time is a controllable parameter. By selecting a longer TR, for instance, we allow more time for the net magnetization vector to recover before we flip it back to the transverse plane for measurement. A longer TR, because it increases the amount of signal that can potentially be detected, will also result in a higher signal-to-noise ratio, with higher image quality.


Fig. 33A.7 Repetition time. This pulse sequence diagram demonstrates the concept of repetition time (TR), which is the time interval between two sequential radio frequency pulses.

(Reprinted with permission from Hashemi, R.H., Bradley, W.G., Lasanti, C.J., 2004. MRI—The Basics. 2nd ed. Lippincott Williams & Wilkins.)

As described earlier, the other process that begins after the initial radiofrequency pulse is turned off is the decrease of net horizontal or transverse magnetization, owing to dephasing of the proton spins (T2 relaxation). Time to echo (TE) refers to the time we wait until we measure the magnitude of the remaining transverse magnetization. TE, just like TR, is a parameter controlled by the operator. If we use a longer TE, tissues with significantly different T2 values (i.e., different rates of loss of transverse magnetization component) will show more difference in the measured signal intensity (transverse magnetization vector size) when the signals are collected. However, there is a tradeoff. If the TE is set too high, the signal-to-noise ratio of the resulting image will drop to a level that is too low, resulting in poor image quality.

Tissue Contrast (T1, T2, and Proton Density Weighting)

By using various TR and TE values, it is possible to increase (or decrease) the contrast between different tissues in an MR image. Achieving this contrast may be based on either the T1 or the T2 properties of the tissues in conjunction with their proton density. Selecting a long TR value reduces the T1 contrast between tissues (Fig. 33A.8). Thus, if we wait long enough before applying the second 90-degree pulse, we allow enough time for all tissues to recover most of their longitudinal or vertical magnetization. Because T1 is relatively short, even for tissues with the longest T1, this is possible without resulting in excessively long scan times. Since after a long TR, the longitudinally oriented net magnetization vectors of separate tissue types are all of similar magnitudes prior to being flipped into the transverse plane by the second pulse, a long TR will result in little T1 tissue contrast. Conversely, by selecting a short TR value, there will be significant variation in the extent to which tissues with different T1 relaxation times will have recovered their longitudinal magnetization prior to being flipped by the second 90-degree pulse (see Fig. 33A.8). Therefore, with a short TR, the second pulse will flip magnetization vectors of different magnitudes into the transverse plane for measurement, resulting in more T1 contrast between the tissues.

During T2 relaxation in the transverse plane, selecting a short TE will give higher measured signal intensities (as a short TE will not allow enough time for significant dephasing, i.e., transverse magnetization loss), but tissues with different T2 relaxation times will not show much contrast (Fig. 33A.9). This is because by selecting a short time until measurement (short TE) we do not allow significant T2-related magnitude differences to develop. If we select longer TE values, tissues with different T2 relaxation times will have time to lose different amounts of transverse magnetization, and therefore by the time of signal measurement, different signal intensities will be measured from these different tissues (see Fig. 33A.9). This is referred to as T2 contrast.

Based on the described considerations, selecting TR and TE values that are both short will increase the T1 contrast between tissues, referred to as T1 weighting. Selecting long TR and long TE values will cause increased T2 contrast between tissues, referred to as T2 weighting.

On T1-weighted images, substances with a longer T1 relaxation time (such as water) will be darker. This is because the short TR does not allow as much longitudinal magnetization to recover, so the vector flipped to the transverse plane by the second 90-degree pulse will be smaller with a lower resulting signal strength. Conversely, tissues with shorter T1 relaxation times (such as fat or some mucinous materials) will be brighter on T1-weighted images, as they recover more longitudinal magnetization prior to their proton spins being flipped into the transverse plane by the second 90-degree pulse (Fig. 33A.10). Among many other applications of T1-weighted images, they allow for evaluation of BBB breakdown: areas with abnormally permeable BBB show increased signal after the intravenous administration of gadolinium. Gadolinium administration is contraindicated in pregnancy. Breast-feeding immediately after receiving gadolinium is generally regarded to be safe (Chen et al., 2008). Renally impaired patients are susceptible to an uncommon but serious adverse reaction to gadolinium, nephrogenic systemic fibrosis (Marckmann et al., 2006).

On T2-weighted images, substances with longer T2 relaxation times (e.g., water) will be brighter because they will not have lost as much transverse magnetization magnitude by the time the signal is measured (Fig. 33A.11). The T1 and T2 signal characteristics of various tissues or substances found in neuroimaging are listed in Table 33A.1.

Table 33A.1 MRI Signal Intensity of Some Substances Found in Neuroimaging

  T1-Weighted Image T2-Weighted Image
Air ↓ ↓ ↓ ↓ ↓ ↓ ↓ ↓
Free water/CSF ↓ ↓ ↓ ↑ ↑ ↑
Fat ↑ ↑ ↑
Cortical bone ↓ ↓ ↓ ↓ ↓ ↓
Bone marrow (fat) ↑ ↑
Edema ↑ ↑
Calcification ↓ (Heavy amounts of Ca++)
↑ (Little Ca++, some Fe+++)
Mucinous material
Gray matter Lower than in T2-WI  
White matter Higher than in T2-WI  
Muscle Similar to gray matter Similar to gray matter
Blood products:    
• Oxyhemoglobin Similar to background
• Deoxyhemoglobin
• Intracellular methemoglobin ↑ ↑
• Extracellular methemoglobin ↑ ↑ ↑ ↑
• Hemosiderin ↓ ↓ ↓

CSF, Cerebrospinal fluid; MRI, magnetic resonance imaging; T2-WI, T2-weighted image.

What happens if we select long TR and short TE values? With the longer TR, the T1 differences between the tissues diminish, whereas the short TE does not allow much T2 contrast to develop. The signal intensity obtained from the various tissues, therefore, will mostly depend on their relative proton densities. Tissues having more proton density, and thereby larger net magnetization vectors, will have greater signal intensity. This set of imaging parameters is referred to as proton density (PD) weighting.

Magnetic Resonance Image Reconstruction

To construct an MR image, a slice of the imaged body part is selected, then the signal coming from each of the voxels making up the given slice is measured. Slice selection is achieved by setting the external magnetic field to vary linearly along one of the three principal axes perpendicular to the axial, sagittal, and coronal planes of the subject being imaged. As a result, protons within the slice to be imaged will precess at a Larmor frequency different from the Larmor frequency within all other imaging planes perpendicular to the axis along which the magnetic field gradient is applied. The Larmor frequency is the natural precession frequency of protons within a magnetic field of a given strength and is calculated simply as the product of the magnetic field, B0, and the gyromagnetic ratio, gamma. The precession frequency of a hydrogen proton is therefore directly proportional to the strength of the applied magnetic field. The gyromagnetic ratio for any given nucleus is a constant, with a value for hydrogen protons of 42.58 MHz/T. In slices at lower magnetic strengths of the gradient, the protons precess more slowly, whereas in slices at higher magnetic field strengths, the protons precess more quickly. Based on the property of nuclear magnetic resonance, the applied radiofrequency pulse (which flips the magnetization vector to the transverse plane) will stimulate only those protons with a precession frequency that matches the frequency of the applied radiofrequency pulse. By selecting the frequency of the stimulating radiofrequency pulse during the application of the slice selection gradient, we can choose which protons (those with a specific Larmor frequency) to stimulate (“make resonate”), and thereby we can select which slice of the body to image (Fig. 33A.12).

After excitation of the slice to be imaged, using the slice selection gradient, the spatial coordinates of each voxel within the slice must be encoded to determine how much signal is coming from each voxel of that slice. This is achieved by means of two additional gradients that are orthogonal to each other within the imaging plane, known as the frequency encoding gradient and the phase encoding gradient. The phase encoding gradient briefly alters the precession frequency of the protons along the axis to which it is applied, thereby changing the relative phases of the precessing protons along this in-plane axis. The frequency encoding gradient, applied orthogonally to the phase encoding gradient within the imaging plane, alters the precession frequency of the protons along the axis to which it is applied, during the acquisition of the MRI signal. As a result of these encoding steps, each voxel will have its own unique frequency and its own unique phase shift, which upon repeating the acquisition with several incremental changes in the phase encoding gradient, will allow for deduction of the spatial localization of different intensity values for each voxel using a mathematical algorithm known as a Fourier transform. Phase encoding takes time; it has to be performed for each row of voxels in the image along the phase encoding axis. Therefore, the higher the resolution of the image along the phase encoding axis, the longer the time required to acquire the image for that slice of tissue.

In the online version of this chapter (available at, there is a discussion of the nature and application of the following MRI sequences or techniques: spin echo and fast (turbo) spin echo; gradient-recalled echo (GRE) sequences, partial flip angle; inversion recovery sequences (FLAIR, STIR); fat saturation; echoplanar imaging; diffusion-weighted magnetic resonance imaging (DWI); perfusion-weighted magnetic resonance imaging (PWI); susceptibility-weighted imaging (SWI); diffusion tensor imaging (DTI); and magnetization transfer contrast imaging.

Magnetic Resonance Imaging

Spin Echo and Fast (Turbo) Spin Echo Techniques

Conventional spin echo imaging is time consuming because the individual echoes are obtained one by one, using a unique strength for the phase encoding gradient at each step in the acquisition of a given slice. The signal from each echo is acquired after a time period equal to one repetition time (TR) after the prior echo. During acquisition and digitization of the signal, with each such step, one row of data space (k-space) is filled. To fill the entire data space for one image, this process has to be repeated as many times as the number of phase encoding steps that are needed. To express this time in seconds, the number of phase encoding steps are multiplied by the TR. Distinct from the conventional spin echo technique, in fast (turbo) spin echo imaging (FSE), within each TR period, multiple echoes at various TE values are obtained, and a new phase encoding step is applied before each of these echoes. The number of echoes obtained for the encoding of each line of k-space in the FSE technique is referred to as the echo train length. Each echo will fill a new line within the k-space data set. Therefore, instead of filling just one line with each TR, multiple lines are filled, and the data space acquisition is completed much more quickly. It is important to realize that even though only a single TE is typically displayed on the MRI technician’s imaging console (this is sometimes referred to as effective TE) during acquisition of FSE images, multiple TE times are actually used. The obvious advantage of fast spin echo imaging is that by filling up k-space much more quickly, the scan time is significantly reduced. This improves image quality by increasing the signal-to-noise ratio. The increased signal, however, may at times also be a disadvantage (e.g., identifying a PV hyperintense lesion adjacent to brighter CSF).

Gradient-Recalled Echo Sequences, Partial Flip Angle

As described earlier, in spin echo imaging, the 90-degree pulse flips the longitudinal magnetization vector into the horizontal plane. After this pulse, the transverse magnetization starts to decay as a result of dephasing. This would in theory result in a decrease of signal by the time (TE) the signal is read by the receiver coils. To prevent this, at a time point equal to one-half of the echo time (TE/2) a 180-degree refocusing pulse is applied that will result in reversal of the directions in which the individual precessing protons are dephasing, so that at a time point equal to TE they will once again be in phase, maximizing the acquired signal by the receiver coils. Thus a signal can be collected that is close in strength to the original. This method only compensates for the dephasing caused by magnetic field inhomogeneities, not for the loss of signal caused by spin-spin interactions, so the recorded signal will not be as large as the original.

In GRE, or gradient echo imaging, instead of “letting” the transverse magnetization dephase and then using the 180-degree refocusing pulse to rephase, a dephasing-refocusing gradient is applied. This gradient will initially dephase the spins of the transverse magnetization. This is followed by the refocusing component of the gradient, which will rephase them at time TE as a readable echo at the receiver coils. Because of greater spin dephasing, GRE is more susceptible to local magnetic field inhomogeneities. This may cause increased artifacts within and near interfaces between tissues with significantly different degrees of magnetic susceptibility, such as at bone/soft tissue or air/bone/brain interfaces near the ethmoid sinuses and medial temporal lobes. However, it is very useful when looking specifically for pathology involving tissue components or deposits exhibiting significant paramagnetism. For example, in the case of chronic hemorrhage, the iron in hemosiderin causes magnetic susceptibility artifact by distorting the magnetic field, resulting in very dark signal voids with an apparent size greater than the spatial extent of the iron deposition, thereby increasing sensitivity for such lesions on the specific pulse sequences designed to maximize this effect. Such pulse sequences include 3D spoiled gradient echo, T2* (pronounced T2-star) and SWI techniques. T2* imaging, in which signal is obtained from transversely magnetized precessing protons without a preceding echo, allows for the detection of hemorrhage as well as deoxyhemoglobin, as in the blood oxygen level dependent (BOLD) effect used to assess relative brain perfusion levels in functional MRI.

Another term that should be explained in conjunction with gradient echo imaging is the partial flip angle. Instead of applying a 90-degree pulse to flip the entire magnetization vector into the horizontal plane, a pulse is used that only partially flips the vector, at a smaller angle. As a result, only a component of the magnetization vector will be in the horizontal plane after application of the excitation pulse. Utilizing a smaller flip angle allows use of a shorter TR, since there will already be a significant longitudinal component of the net magnetization vector after excitation, requiring less time for sufficient recovery of longitudinal magnetization prior to the next excitation pulse.

The T1-weighted signal generated by a tissue in a GRE sequence can be optimized for any given TR by varying the flip angle according to a mathematical relationship known as the Ernst equation. The optimal flip angle for a given tissue at a particular TR is thus known as the Ernst angle.

Use of shorter longitudinal relaxation times in gradient echo imaging has the obvious advantage of decreasing scan time. By changing the flip angle (which, just like TR and TE is an operator-controlled parameter) the tissue contrast may be manipulated. Selecting a small flip angle in conjunction with a sufficiently long TR will decrease the T1 weighting of the image, as the longitudinal magnetization will be nearly maximized for all tissues. This effect is similar to that for a conventional spin echo sequence, when selecting a long TR allows the longitudinal magnetization to recover more, thereby reducing or eliminating T1 weighting from the resulting image.

The generation of image contrast in GRE imaging is similar to that in spin-echo imaging. One important difference is that T2-weighted images cannot be generated, owing to lack of a refocusing pulse in the GRE technique. Instead, the shorter T2* decay is used to generate T2-like image contrast while minimizing T1 effects. Therefore, T2*-weighted images are obtained using a small flip angle, a long TR, and long TE. A small flip angle in conjunction with a long TR and a short TE will result in proton density weighting, because the T1 and T2* effects upon image contrast are minimized. Selecting a large flip angle together with a short TR and a short TE will result in T1 weighting. Advantages of GRE imaging include speed, less contamination of signal in the slice to be imaged by signal from adjacent slices, and higher spatial resolution. Disadvantages include greater susceptibility to inhomogeneities in the magnetic field such as magnetic susceptibility artifact (although this may also be an advantage, as outlined earlier) and the requirement for higher gradient field strengths. One very useful application of GRE imaging is in volumetric analysis of imaged tissues; the shorter TR and resultant speed allow for rapid data acquisition in three dimensions, which can be used to format and display images in any plane.

Inversion Recovery Sequences (FLAIR, STIR)

For better detection and visualization of abnormalities on MR images, it is often useful to suppress the signal from certain tissues, thereby increasing the contrast between the region of pathology and the surrounding and/or background tissue. Examples of this include visualization of hyperintense lesions adjacent to bright CSF spaces on T2-weighted images, or whenever there is a need to eliminate the hyperintense signal coming from fatty background.

Inversion recovery techniques use a unique pulse sequence to avoid signal detection from the selected tissues (fat or CSF). Initially, a 180-degree radiofrequency pulse is applied. This will flip the longitudinal magnetization vectors of all tissues by 180-degrees, so that the vectors will point downward (south). Next, the flipped vectors are allowed to start recovering according to their respective T1 times. As the downward-pointing vectors recover, they become progressively smaller, eventually reaching zero magnitude, and from that point they start growing and pointing upward (north). Without interference, they recover the original longitudinal magnetization. However, during the process of recovery, after a time period referred to as inversion time (TI), a 90-degree pulse is applied. This will flip the longitudinal vectors to the transverse plane, where signal detection occurs. The amount of magnetization flipped by this pulse depends on how far the longitudinal recovery has been allowed to proceed. If the 90-degree pulse is applied when a given tissue’s vector happens to be zero (this is the so-called null point), no magnetization will be flipped from that tissue to the transverse plane, and therefore no signal will be detected from that tissue. Different tissues recover their longitudinal magnetization at different rates according to their specific T1 times. Knowing a given tissue’s T1 time, we can calculate when it will reach the null point (when its longitudinal magnetization is zero), and if we apply the 90-degree pulse at that point, we will not detect any signal from that particular tissue. The inversion time is linearly dependent upon a given tissue’s T1 value, being calculated as 0.69 multiplied by the T1 value. In the FLAIR (fluid-attenuated inversion recovery) sequence, the inversion time (when the 90-degree pulse is applied) occurs when the magnetization vector for the CSF is at the null point, so no signal will be detected from the CSF (Fig. 33A.13). In FLAIR images, the dark CSF is in sharp contrast with the hyperintensity of PV lesions, allowing their better identification. In STIR (short TI, or tau inversion, recovery) imaging, which is a fat-suppression technique, the methodology is essentially the same as for FLAIR. However, instead of CSF, the signal from fat is nulled. The TI for the STIR technique is set to 0.69 times the T1 of fat, which results in application of the final 90-degree pulse when the fat tissue’s magnetization is at the null point, so no signal from fat will be detected.

Diffusion-Weighted Magnetic Resonance Imaging

Diffusion of water molecules within tissues has a random molecular (Brownian) motion, which varies in a tissue- and pathology-dependent manner. It may have a directional preference in some tissues; for instance, there is greater diffusion in the longitudinal than in the transverse plane of an axon. Water diffusion may occur more rapidly in aqueous compartments such as CSF, relative to water that is largely intracellular, as in regions of cytotoxic edema secondary to brain ischemia or water present in fluid compartments with high viscosity, such as abscesses or epidermoid cysts. DWI is an imaging technique that is able to differentiate areas of low from high diffusion. The imaging sequence used for this purpose is a T2-weighted sequence, with the addition of transient gradients applied before TE. The spin-echo sequence can be conventional, or more commonly, a single-shot spin-echo echoplanar imaging sequence that allows for much faster imaging. The purpose of the gradients is to sensitize the pulse sequence to diffusion occurring during the time interval between their application. In tissues where more diffusion occurred during application of the gradient (such as in normal tissues), the diffusion causes dephasing of transverse magnetization, resulting in signal loss and, therefore, a darker appearance on the image. In areas with less diffusion (for example in acutely ischemic brain areas), no significant dephasing or signal change occurs. Therefore, the detected signal is higher, and these areas appear bright on the image.

The degree of the applied diffusion encoding gradient is referred to as the B value. In a regular conventional T2 or FLAIR image, the B value is zero (i.e., no gradient). As the B value is increased by the gradient being stronger, the diffusion of the water molecules will cause more and more dephasing and signal loss. As a result, if the B value is high enough, as in DWI, the areas of higher diffusion rates, such as CSF and normal brain tissue, will be dark due to the dephasing and signal loss related to water diffusion. In contrast, ischemic areas with little or no water molecule diffusion will appear bright because they lack dephasing and signal loss. In imaging protocols where more T2 weighting (longer TE values) and smaller B values are used, areas with long T2 values may appear relatively bright in the diffusion-weighted images, despite their considerable diffusion. This phenomenon is referred to as T2 shine-through, and it is due to the low applied B value, which means a weaker diffusion gradient and less diffusion weighting. This shine-through can be decreased by applying a stronger diffusion gradient, leading to higher B values and more diffusion weighting.

Based on the differences in the change of signal intensity in different areas at different applied B values, it is possible to calculate the apparent diffusion coefficient (ADC) in various areas/tissues in the image. The term apparent is used because in a tissue there are other factors besides this coefficient that contribute to signal loss, including patient motion and blood flow. The higher the diffusion rate, the higher the ADC value of the given tissue, and the brighter it will appear on the ADC image or map. As an example, CSF, where the diffusion is highest, will be bright on the ADC map, whereas areas of little (restricted) diffusion, such as ischemic areas, will be dark.

One of the most obvious practical uses of DWI is the delineation of acutely ischemic areas, which appear bright against a dark background in diffusion-weighted images and dark on the corresponding ADC maps. According to the most appealing theory, the reason for restricted diffusion in acutely ischemic brain tissue is the evolving cytotoxic edema (cellular swelling), which decreases the relative size of the extracellular space, thereby limiting water diffusion.

Although in neurological practice, the term restricted diffusion usually refers to cerebral ischemia, and this imaging modality remains most important for acute stroke imaging, it is to be noted that there are other abnormalities that also restrict diffusion and appear bright on diffusion-weighted images. Examples include abscesses, hypercellular tumors such as lymphoma, some meningiomas, epidermoid cysts, aggressive demyelinating disease, and proteinaceous material, such as in association with sinusitis.

Perfusion-Weighted Magnetic Resonance Imaging

Perfusion-weighted imaging utilizes MRI sequences that are able to generate different signal intensities between tissues with different degrees of perfusion. Although there are techniques (like spin-labeled perfusion imaging) that provide information about tissue perfusion without injecting contrast material, the most common technique uses a rapid bolus of paramagnetic contrast agent (gadolinium) which, while passing through the tissues, causes distortion of the magnetic field and signal loss in the applied gradient echo or echo planar image. This signal loss only occurs in tissues that are perfused, whereas nonperfused regions do not have such signal loss, or in cases of decreased but not absent perfusion, the signal loss is not as prominent as seen in the healthy tissue. When the selected slice is imaged repeatedly in rapid succession, parameters related to perfusion (e.g., relative cerebral blood volume [rCBV], time to peak signal loss [TTP], mean transit time of the contrast bolus [MTT]) can be calculated for each voxel within the slice being imaged. Estimates of cerebral blood flow (CBF) can be calculated for each voxel as well.

The main clinical application of PWI is in the setting of acute stroke, primarily for visualization of tissue at risk, the ischemic penumbra. When used in conjunction with diffusion-weighted images, which delineate the acutely infarcting area, it is frequently seen that perfusion-weighted images reveal a more extensive area, beyond the extent of the zone of infarction, that exhibits decreased or absent perfusion. This is the ischemic penumbra, tissue at risk that is potentially salvageable, prompting use of thrombolytic therapy. If the perfusion deficit appears the same as the zone of restricted diffusion (area in the process of infarction), the chance for saving tissue is likely to be lower than that for an ischemic infarction exhibiting a significant perfusion-diffusion mismatch.

Susceptibility-Weighted Imaging

As described earlier, factors that distort magnetic field homogeneity, such as paramagnetic or ferromagnetic substances, cause local signal loss. Signal loss occurs because in the altered local magnetic field, protons will precess with different frequencies, resulting in dephasing and thus decreasing the net magnetization vector that translates into a detectable signal. Gradient echo images are especially sensitive to magnetic field distortions, which appear as areas of decreased signal due to the magnetic susceptibility artifact.

SWI (Haacke et al., 2009; Mittal et al., 2009) uses a high spatial resolution 3D gradient echo imaging sequence. The contrast achieved by this sequence distinguishes the magnetic susceptibility difference between oxygenated and deoxygenated hemoglobin. Since the applied phase postprocessing sequence accentuates the paramagnetic properties of deoxyhemoglobin and blood degradation products such as intracellular methemoglobin and hemosiderin, this technique is very sensitive for intravascular venous deoxygenated blood as well as extravascular blood products. It has been used for evaluation of venous structures, hence the earlier name high-resolution blood oxygen level–dependent venography, but the clinical application is now much broader. Its exquisite sensitivity for blood degradation products makes this technique very useful when evaluating any lesion (e.g., stroke, AVM, cavernoma or neoplasm) for associated hemorrhage (Fig. 33A.15). It is also used for imaging microbleeds associated with traumatic brain injury, diffuse axonal injury, or cerebral amyloid angiopathy.

Diffusion Tensor Imaging

Diffusion tensor imaging is a more advanced type of diffusion imaging capable of quantifying anisotropy of diffusion in white matter. Diffusion is isotropic when it occurs with the same intensity in all directions. It is anisotropic when it occurs preferentially in one direction, as along the longitudinal axis of axons. For this reason, DTI finds its greatest current application in MRI examinations of the white matter. As opposed to characterizing diffusion within each voxel with just a single apparent diffusion coefficient, as in DWI, in DTI intravoxel diffusion is measured along three, six, or more gradient directions. The measured values and their directions are called eigenvectors. The vector that corresponds to the principal direction of diffusion (the direction in which diffusion is greatest in magnitude) is called the principal eigenvector. In normal white matter, diffusion anisotropy is high because diffusion is greatest parallel to the course of the nerve fiber tracts. Therefore, the principal eigenvector delineates the course of a given nerve fiber pathway. Diffusion tensor images can be displayed as maps of the principal eigenvectors which will show the direction/course of the given white matter tract (tractography). These images can also be color coded, allowing for more spectacular visualization of nerve fiber tracts (Fig. 33A.16). Any disruption of a given nerve fiber tract (e.g., MS, trauma, gliosis) will reduce anisotropy, and the disruption of the white matter tract can be visualized. Tensor imaging/tractography is useful in imaging of degenerating white matter tracts and also in surgical resection planning, when the anatomical relationship of the resectable lesion and the adjacent fiber tracts has to be evaluated to avoid or reduce surgical injury to critical pathways.

Magnetization Transfer Contrast Imaging

As the name indicates, magnetization transfer contrast imaging is a technique that produces increased contrast within an MR image, specifically on T1-weighted gadolinium-enhanced images and in magnetic resonance angiography (Henkelman et al., 2001). In water, hydrogen atoms are relatively loosely bound to oxygen atoms, and they move frequently between them, binding to one oxygen atom then switching to another. In other tissues (e.g., lipids, proteins), the hydrogen atoms are more tightly bound and tend to stay in one place for longer periods of time. Nevertheless, it does happen that a “bound” hydrogen in lipid or protein is exchanged with a “more free” hydrogen from water. In magnetization transfer imaging, at the beginning of the sequence a radiofrequency pulse is applied that saturates the bound protons in lipids and proteins but does not affect the free protons in water. In regions where magnetization transfer (i.e., exchange of saturated protons with free protons) occurs, the saturated protons will decrease the signal obtained from the imaged free protons. The more frequently this magnetization transfer occurs, the less signal is obtained from the region and the darker the region will be in the image. Magnetization transfer happens more frequently in the white matter, resulting in signal loss, and therefore on magnetization transfer images, the white matter appears darker. The CSF on the other hand, where magnetization transfer does not occur, does not lose signal. Magnetization transfer is minimal in blood because of the high amount of free water protons.

This technique is useful when gadolinium-enhanced T1-weighted images are obtained, because enhancing lesions stand out better against the darker background of the more hypointense white matter. In fact, applying a magnetization transfer sequence to single-dose gadolinium-enhanced T1-weighted images results in contrast enhancement intensity comparable to giving a double dose of gadolinium. This sequence is also used in time-of-flight magnetic resonance angiography. There is no signal change in the blood, but the background tissue becomes darker, so the imaged blood vessels stand out better, and smaller branches are better visualized. This benefit comes at the expense of a significantly prolonged scan time, because it takes additional time to apply the magnetization transfer pulse.

Another application of magnetization transfer imaging is in the assessment of “normal-appearing” tissues that in fact contain abnormalities, albeit not visible on conventional MR pulse sequences. By selecting a region of interest (ROI, essentially a quadrilateral that is selected to enclose the tissue of interest within an image) corresponding to the “normal-appearing” tissue and calculating the degree to which magnetization transfer occurs within each voxel of the ROI, a histogram plot can be generated. On such magnetization transfer ratio (MTR) histograms, tissues with no apparent lesional signal on conventional images, such as the “normal-appearing white matter” of MS, may exhibit a decreased peak height. Such histograms in MS patients may also exhibit a larger proportion of voxels with low MTR values than normal tissues, reflecting a microscopic and macroscopic lesion load that is otherwise undetectable by conventional imaging techniques.

Structural Neuroimaging in the Clinical Practice of Neurology

Brain Diseases*

Brain Tumors

Epidemiology, pathology, etiology, and management of cancer in the nervous system are discussed in Chapter 52A, Chapter 52B, Chapter 52C, Chapter 52D, Chapter 52E, Chapter 52F, Chapter 52G . From the standpoint of structural neuroimaging, a useful anatomical classification distinguishes two main groups: intraaxial and extraaxial tumors. Intraaxial tumors are within the brain parenchyma, extraaxial tumors are outside the brain parenchyma (involving the meninges or, less commonly, the ventricular system). Intraaxial tumors are usually infiltrative with poorly defined margins. Conversely, extraaxial tumors, even though they often compress or displace the adjacent brain, are usually demarcated by a cerebrospinal (CSF) cleft or another tissue interface between tumor and brain parenchyma. For differential diagnostic purposes, intraaxial primary brain neoplasms can be further divided into the anatomical subgroups of supratentorial and infratentorial tumors (Table 33A.2).

For evaluation of brain tumors, the structural imaging modality of choice is MRI. Due to their gradual expansion and often infiltrative nature, most brain tumors are already visible on MRI by the time patients become symptomatic. Exceptions to this rule are tumors that tend to involve the cortex or corticomedullary junction, such as small oligodendrogliomas or metastases, which may cause seizures early, even before being clearly visible on noncontrast MRI. Meningeal involvement is also often symptomatic, for instance by causing headaches and confusion, but may not be appreciated on noncontrast images. Higher magnetic field strength (e.g., a 3-T scanner) and contrast administration (in double or triple dose if necessary) can improve detection of small or clinically silent neoplastic lesions.

Neuroimaging is particularly useful in the assessment of brain tumors. Unlike destructive lesions such as ischemic strokes, brain tumors often cause clinical manifestations that are difficult to interpret. Sometimes the clinical presentation may provide clues to localization—for example, a seizure is suggestive of an intraaxial tumor, whereas cranial nerve involvement tends to signal an extraaxial pathology. But edema, mass effect, obstructive hydrocephalus, and elevated intracranial pressure (ICP) can give rise to nonspecific symptoms (e.g., headache, visual disturbance, altered mental status), and false localizing signs may also appear, such as oculomotor or abducens nerve compression due to an expanding intraaxial mass.

Neoplastic tissues most commonly prolong the T1 and T2 relaxation times, appearing hypointense on T1 and hyperintense on T2-weighted images, but different tumors differ in this property, facilitating tumor identification on MRI. MRI is also very sensitive for detection of other pathologic changes that can be associated with tumors, such as calcification, hemorrhage, necrosis, and edema. The structural detail provided by MRI is useful for assessing involved structures and determining the number and macroscopic extent of the neoplasms, thereby guiding surgical planning or other treatment modalities.

Intraaxial Primary Brain Tumors

Pilocytic Astrocytomas

Pilocytic astrocytomas have two major groups: juvenile and adult. These tumors are classified as WHO grade I. Juvenile pilocytic astrocytomas are the most common posterior fossa tumors in children. The most common locations are the cerebellum, at the fourth ventricle, third ventricle, temporal lobe, optic chiasm, and hypothalamus (Koeller and Rushing, 2004). The appearance is often lobulated, and the lesion appears well demarcated on MRI. Hemorrhage and necrosis are uncommon. Areas of calcification may be present. The tumor usually exhibits solid as well as cystic components, with or without a mural nodule. The adult form is usually well circumscribed, often calcified, and typically exhibits a large cyst with a mural nodule. On MRI, the solid portions of the tumor are iso- to hypointense on T1 and iso- to hyperintense on T2-weighted images (Arai et al., 2006). The cystic component usually exhibits CSF signal characteristics. The associated edema and mass effect is usually mild, sometimes moderate. With gadolinium, the solid components (including the mural nodule) enhance intensely, but not the cyst, which rarely may show rim enhancement.

Pleomorphic Xanthoastrocytoma

Pleomorphic xanthoastrocytoma is a rare variant of astrocytic tumors. It is thought to arise from the subpial astrocytes and typically affects the cerebral cortex and adjacent meninges and may cause erosion of the skull. The most common location is the temporal lobe. It is classified as WHO grade II. It usually occurs in the second and third decades of life, and patients often present with seizures. On MRI (Tien et al., 1992) usually a well-circumscribed cystic mass appears in a superficial cortical location. A solid portion or mural nodule is often seen, and the differential diagnosis includes pilocytic astrocytoma and ganglioglioma. The signal characteristics are hypointense or mixed on T1, and hyperintense or mixed on T2-weighted images. With contrast, the solid portions and sometimes the adjacent meninges enhance. Calcification may be present. There is mild or no mass effect associated with this tumor.

Low-Grade Astrocytomas

Fibrillary astrocytomas, also termed diffuse astrocytomas, represent approximately 10% of all gliomas. Low-grade (WHO grade I and II) astrocytomas belong to this group (Figs. 33A.17 and 33A.18). These are well-differentiated tumors, usually arising from the fibrillary astrocytes of the white matter. Even though imaging may show a fairly well-defined boundary, these tumors are infiltrative and usually spread beyond their macroscopic border. Two-thirds of cases are supratentorial. A subgroup of these astrocytomas involves specific regions such as the optic nerves/tracts or the brainstem.

Low-grade astrocytomas are iso- or hypointense on T1-weighted images and hyperintense on T2-weighted images. Tissue expansion may be seen, and mass effect (if present) is generally modest. There is little to no associated edema. Fibrillary grade I astrocytomas do not enhance; grade II tumors may exhibit enhancement. The appearance of enhancement in a previously nonenhancing tumor is a worrisome sign of malignant transformation, often due to anaplastic astrocytoma.