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Figure 3.50 The parallel protons will align with B0, but will be out of phase.


Figure 3.51A The protons will all be in phase, forming a net magnetic vector (NMV).

they are aggregately termed the net magnetic vector (NMV) (Figure 3.51A). Then, the NMV will be pushed from a low-energy parallel alignment to a higher-energy transverse alignment (Figure 3.51B). They will remain in this state as long as external RF energy is applied.

Once the external RF energy is terminated, the NMV will begin to lose the absorbed energy (Figure 3.52) and simultaneously go through two separate processes. First, the NMV will begin to relax from the excited state to the preferred longitudinal state. As it does so, it gives up its accumulated energy to the surrounding tissue lattice. This process is called T1 recovery or spin lattice. T1 time is defined as the time is taken to recover 63% of the longitudinal magnetization (Figure 3.53). The second process that affects the NMV involves the individual protons within its bulk. Once the external RF energy is terminated, the individual protons begin to dephase and lose step with one another. This process is driven by the interaction and exchange of energy between the individual protons and is called T2 decay. T2 decay is defined as the time taken for 37% of the transverse magnetization to be lost (Figure 3.54). The rate at which both these processes occur is influenced by the strength of the magnetic field and the biological state of the imaged tissue. The latter plays a critical role when imaging preserved, dried, and/or decaying tissues.

During the T1 and T2 processes, the NMV, an aggregate of multiple small magnetic fields, is precessing on its axis. This magnetic field, though small, is precessing at 8-127 MHz


Figure 3.51B The radiofrequency will be absorbed, and in doing so the NMV will be pushed into the higher-energy transverse alignment.


Figure 3.51B The radiofrequency will be absorbed, and in doing so the NMV will be pushed into the higher-energy transverse alignment.

Figure 3.52 once the RF is terminated, the protons begin to lose the absorbed energy and return to B0, the longitudinal magnetization.

depending on field strength, 0.2-3.0 T accordingly. According to Faraday's laws of induction, a magnetic field of changing intensity perpendicular to a wire will induce a voltage along the length of that wire. MRI exploits the fact by placing a special wire conductor, called a coil, perpendicular to the precessing NMV. There are coils of many shapes and sizes, and it is best to choose the smallest coil possible for the body part to be imaged. Since the amount of voltage induced depends on both the strength and rate of change of the magnetic field, we can see that even a weak magnetic field is capable of inducing a measurable current when precessing between 8 and 127 million times per second. The induced current, similar to the NMV itself, is not a single entity but composed of the sum of its parts and is of the same frequency as the precession of the NMV. Careful analysis of the signal will yield valuable information about the protons that created it. As discussed earlier, the rate at which T1 and T2 occur is dependent on the nature of the tissues. By manipulating the excitation and relaxation process, we can, in essence, interrogate the protons and extract valuable information regarding the quantity and binding characteristics of water molecules.


Figure 3.53 The T1 time signifies the loss of energy to the surrounding tissue. The process is defined as the amount of time taken to recover 63% of the longitudinal magnetization.


Figure 3.53 The T1 time signifies the loss of energy to the surrounding tissue. The process is defined as the amount of time taken to recover 63% of the longitudinal magnetization.

Figure 3.54 T2 decay is defined as the time taken for 37% of the transverse magnetization to be lost.

The information gained through this manipulation, though valuable, is limited to two dimensions: frequency and amplitude. To create an image, we need to add information regarding the precise location of each component of the signal within the sample. This is done through a process called spatial localization. Although a detailed understanding of spatial localization is not necessary for this text, the basic concept should be mentioned.

The entire basis of clinical MRI is that protons will precess in a magnetic field. The rate of precession is extremely precise, and for hydrogen it is equal to 42.57 MHz per tesla. We can vary the rate of precession by varying the magnetic field, and by doing so we can vary the frequency of the signal induced in the coil. If in addition to the main magnetic field, we apply changeable magnetic fields in the X, Y, and Z planes, then we have the ability to alter the precession in any direction we choose. These additional components are called gradient coils and are really nothing more than electromagnets arranged around the bore (tunnel) of the main magnet. Though very complex in application, the concept is simple. Apply a changing electromagnetic field by altering power to the gradient in any one direction, and you will alter the precession of the individual protons along that axis. Record the values, and repeat the process in another plane. By repeating this process thousands of times per second, we can not only determine the T1 and T2 properties of the imaged tissues but also their precise anatomical location within the sample. The T1 and T2 properties are important because they tell us the type of tissue and its general state of health. The location is important because it allows us to image not in a single slice, like early CT scanners, but rather across a large volume of tissue that the computer can segment into slices for easy viewing. The added benefit of MRI is that since the image is created not by the attenuation of an x-ray beam but rather by the concentration and distribution of water molecules, the image is extremely sensitive to subtle tissue changes that would otherwise be lost to other modalities.

The disadvantage of this increased sensitivity is that it comes with a much more complex series of user-defined parameters for obtaining images. Whereas x-ray is limited to a few basics and CT a few more, MRI has dozens of user-selectable parameters that create hundreds of thousands of possible combinations. With MRI in widespread clinical use since the mid- to late 1980s, the parameters are well understood for clinical human imaging. In the end, we are really trying to obtain an image that is optimized for either the T1 properties of a tissue, the T2 properties of the tissue, or the overall density of protons contained in the tissue. We do this by manipulating the parameters within long-established guidelines.

Table 3.1 signal Characteristics of Tissues with spin Echo T1 and T2 Imaging

Hyperintense on T1

Hyperintense on T2

Subacute blood



Highly proteinaceous fluids Cholesterol

Gadolinium enhancement

Water CSF

Subacute hematoma Inflammation

Highly proteinaceous fluids Most lesions

The basic schematic diagram of how the excitation and relaxation processes are controlled along with the functions of the gradients is called a pulse sequence. Though complex in application, the pulse sequence is nothing more than a preset series of events that the computer uses to produce an image. The pulse sequence has several components that are controllable by the operator, and a few of these basic parameters are applicable here.

The few parameters we will mention in this text will be those with the greatest impact on the images we wish to obtain. TE (echo time) is defined as the time between the RF excitation pulse and the time we sample the resultant signal. The longer the TE, the less the signal available, because of the rapid rate of the T1 and T1 processes. TR (repetition time) is defined as the time from one excitation pulse to the next, or how long we wait before repeating the entire process. By altering these values, we can formulate an image with the tissue characteristics we desire to see (Table 3.1). These two parameters have a great impact on the signal characteristics of the resultant image. For example, fat will be bright and water will be dark.

The other parameters we set have little impact on the signal characteristics, but they play a critical role in the overall quality of the image. When discussing image quality, we are discussing a trade-off between image resolution, or detail, and signal-to-noise ratio (SNR). An image can have a very high resolution but a low SNR. Such an image will be of little value because although we have great detail, there is not enough signal to see it. The opposite is also true. A high-SNR image with low resolution, though pretty, will not provide sufficient detail to visualize the intended structures. The key parameters needed to balance SNR and resolution are field of view (FOV), slick thickness, and image matrix. In short, the more protons included in each portion of the image, the greater the SNR. Thick slices and large FOVs accomplish this nicely. Image resolution is more a factor of the displayed image. The more tiny pieces an image is divided into (pixels), the better the detail. But if the pieces are too small or if the images were collected with a small FOV and thin slices, there may not be enough signal to fill the pixels and the resultant image will be very noisy.

Images can be classified as T1 or T2 weighted depending on which process contributed more to the final image. With a T1-weighted image, fat has a higher signal than water (Figure 3.55). High signal appears white on the displayed image, lower signals are generated as shades of gray, and no signal, such as received from air or cortical bone, is black. T1 images provide excellent anatomical detail. On the other hand, a T2-weighted image provide a higher signal from water and a lower signal from fat (Figure 3.56). T2 images are advantageous for demonstrating pathological processes,

Figure 3.55 A T1-weighted image at the level of the cerebellum (A). Note the high signal from the retro-orbital fat.

such as edema, that have high "water" content. However, T2 images don't provide as much anatomical detail as Tl-weighted images. Through careful manipulation of the technical parameters, a skilled operator can create images to enhance or feature almost any disease process.

Because MR primarily provides information on the location of mobile hydrogen within a body, it would seem of little value in the examination of mummified remains. Notman et al. (1986) published a report following the examination of a mummy with MR looking for residual moisture. They stated, "... It appears MRI is unsuitable for the paleopathologic investigation of dehydrated structures." At the time they were correct, as early attempts to utilize MRI to image mummies and mummy tissues realized little success. Three notable exceptions are a modern mummy created in the ancient Egyptian method in 1994 at the University of Maryland by Bob Brier and Ronn Wade (Quigley 1998), a dog mummified to document and analyze the desiccation process by Notman and Aufderheide (1995), and brain tissue removed from the preserved skull of an Indian Bog Mummy and embedded in agar (Notman 1983).

Since that time, we have, on several occasions, been successful in incorporating MRI into the imaging workup protocol for the evaluation of mummified remains. With mummified tissues, it is sometimes necessary to override the standard prescan process and tune the system manually. It is also helpful in some cases to add an isolated source of mobile protons to "load" the coil. This can be accomplished by placing saline IV bags or water bottles in the coil with the body part to be imaged. The mobile protons in the water trick the system into "thinking" that the tissue within the coil is hydrated. Once the system tunes in appropriately, the operator can fine-tune the scanner and obtain satisfactory images even with the bags removed.


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Figure 3.56 A T2-weighted image at approximately the same level as in Figure 3.55. In this case, the high signal is from the aqueous fluid in the eye (A) and the cerebrospinal fluid (B) anterior and lateral to the pons.

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