Apparent Diffusion Coefficient

(ADC) A diffusion coefficient to differentiateT2 shine through effects or artifacts fromreal ischemic lesions. In the human brain, water diffusion is a three-dimensional process that is not truly random because the diffusional motion of water is impeded by natural barriers. These barriers are cell membranes, myelin sheaths, white matterfiber tracts, and protein molecules.
The apparent water diffusion coefficients can be calculated by acquiring two or more images with a different gradient duration and amplitude (b-values). The contrast in the ADC map depends on the spatially distributed diffusion coefficient of the acquired tissues and does not contain T1and T2* values.
The increased sensitivity of diffusion-weighted MRI in detecting acute ischemia is thought to be the result of the water shift intracellularly restricting motion of water protons (cytotoxic edema), whereas the conventional T2 weighted images show signal alteration mostly as a result of vasogenic edema. 
The reduced ADC value also could be the result of decreased temperature in the nonperfused tissues, loss of brain pulsations leading to a decrease in apparent proton motion, increased tissue osmolality associated with ischemia, or a combination of these factors. The lower ADC measurements seen with early ischemia, have not been fully established, however, a lower apparent ADC is a sensitive indicator of early ischemicbrain at a stage when ischemic tissue remains potentially salvageable.
See also Diffusion Weighted Imaging and Diffusion Tensor Tractography.

FLAIRPRO

(IR) Inversion recovery is an MRI technique, which can be incorporated into MR imaging, wherein the nuclear magnetization is inverted at a time on the order of T1 before the regular imaging pulse-gradientsequences. The resulting partial relaxation of the spins in the different structures being imaged can be used to produce an image that depends strongly on T1. This may bring out differences in the appearance of structures with different T1 relaxation times. Note that this does not directly produce an image of T1. T1 in a given region can be calculated from the change in the MR signal from the region due to the inversion pulse compared to the signal with no inversionpulse or an inversion pulse with a differentinversion time. This sequence involves successive 180° and 90° pulses. The inversion recovery sequence is specified in terms of three parameters, inversion time (TI), repetition time (TR) and echo time (TE).
See also Inversion Recovery Sequence and FLAIR.

SPIN RADIOGRAPHERS: MRS

SPIN RADIOGRAPHERS: MRS: ( MRS  /  MRS I -  Magnetic Resonance Spectroscopic Imaging) A method using the NMR  phenomenon to identify the chemical state of various...

MRS

(MRS / MRSI - Magnetic ResonanceSpectroscopic Imaging) A method using theNMR phenomenon to identify the chemical state of various elements without destroying the sample. MRS therefore provides information about the chemical composition of the tissues and the changes in chemical composition, which may occur with disease processes.
Although MRS is primarily employed as a research tool and has yet to achieve widespread acceptance in routine clinical practice, there is a growing realization that a noninvasive technique, which monitors disease biochemistry can provide important new information for the clinician. 
The underlying principle of MRS is that atomic nuclei are surrounded by a cloud of electrons, which very slightly shield the nucleus from any external magnetic field. As the structure of the electron cloud is specific to an individual molecule or compound, then the magnitude of this screening effect is also a characteristic of the chemical environment of individual nuclei.
In view of the fact that the resonant frequency is proportional to the magnetic field that it experiences, it follows that the resonant frequency will be determined not only by the external applied field, but also by the small field shift generated by the electron cloud. This shift in frequency is called the chemical shift (see alsoChemical Shift). It should be noted that chemical shift is a very small effect, usually expressed in ppm of the main frequency. In order to resolve the different chemical species, it is therefore necessary to achieve very high levels of homogeneity of the main magnetic field B0. Spectra from humans usually require shimmingthe magnet to approximately one part in 100. High resolution spectra of liquid samples demand ahomogeneity of about one part in 1000.
In addition to the effects of factors such as relaxation times that can affect the NMR signal, as seen inmagnetic resonance imaging, effects such as J-modulation or the transfer of magnetization after selective excitation of particular spectral lines can affect the relative strengths of spectral lines.
In the context of human MRS, two nuclei are of particular interest - H-1 and P-31. (PMRS - Proton Magnetic Resonance Spectroscopy) PMRS is mainly employed in studies of the brain where prominent peaks arise from NAA, choline containing compounds, creatine and creatine phosphate, myo-inositol and, if present, lactate; phosphorus 31 MR spectroscopy detects compounds involved in energy metabolism (creatine phosphate, adenosine triphosphate and inorganic phosphate) and certain compounds related to membrane synthesis and degradation. The frequencies of certain lines may also be affected by factors such as the local pH. It is also possible to determine intracellular pH because the inorganic phosphate peak position is pH sensitive.
If the field is uniform over the volume of the sample, "similar" nuclei will contribute a particular frequencycomponent to the detected response signal irrespective of their individual positions in the sample. Since nuclei of different elements resonate at different frequencies, each element in the sample contributes a different frequency component. A chemical analysis can then be conducted by analyzing the MR response signal into its frequency components.
See also Spectroscopy.

Time of Flight Angiography

(TOF) The time of flight angiography is used for the imaging of vessels. Usually the sequence type is a gradient echo sequences with short TR, acquired with slices perpendicular to the direction of blood flow.
The source of diverse flow effects is the difference between the unsaturated and presaturated spins and creates a bright vascular image without the invasive use of contrast media. Flowing blood moves unsaturated spins from outside the slice into the imaging plane. These completely relaxed spins have full equilibrium magnetization and produce (when entering the imaging plane) a much higher signal than stationary spins if a gradient echo sequence is generated. This flow related enhancement is also referred to as entry slice phenomenon, or inflow enhancement.
Performing a presaturation slab on one side parallel to the slice can selectively destroy the MR signal from the in-flowing blood from this side of the slice. This allows the technique to be flow direction sensitive and to separate arteriograms or venograms. When the local magnetization of moving blood is selectively altered in a region, e.g. by selective excitation, it carries the altered magnetization with it when it moves, thus tagging the selected region for times on the order of the relaxation times.
For maximum flow signal, a complete new part of blood has to enter the slice every repetition (TR) period, which makes time of flight angiography sensitive to flow-velocity. The choice of TR and slice thickness should be appropriate to the expected flow-velocities because even small changes in slice thickness influences the performance of the TOF sequence. The use of sequential 2 dimensional Fourier transformation (2DFT) slices, 3DFT slabs, or multiple 3D slabs (chunks) are depending on the coverage required and the range of flow-velocities.
 

FIESTA

(FIESTA) The fast imaging employing steady state acquisition sequence provides images of fluid filled structures with very short acquisition times. The FIESTA sequence uses the T2 steady state contrast mechanism to provide high SNR images with strong signal from fluid tissues while suppressing background tissue for contrast and anatomic detail of small structures. In addition, the ultra short TR and TE enable extremely short acquisition times – shorter than FSE – and the images can be post processed using MIP, volume rendering, or 3D navigator techniques.

Helical CT Evaluation of the Thoracic Aorta: Helical CT Aortography Technique

 Although no standardized techniques exist for helical CT evaluation of the thoracic aorta, and practices vary from institution to institution, certain general principles may be applied to optimize techniques. Variables that must be considered include collimation, pitch, field-of-view, reconstruction increment, amount, rate and timing of contrast administration, distance to be scanned, kVp, mA, and tube rotation time. These variables must be optimized to provide the highest possible scan quality. Occasionally, a compromise among these parameters is required to achieve this goal.

Collimation and Pitch

When a scan protocol is constructed, one must generally strike a compromise between the collimation and the pitch. Recall that pitch reflects the ratio of table speed to slice thickness (collimation) per tube rotation, or:

Pitch =

Table transport distance Collimation

A useful general rule of thumb is that effective slice thickness increases by 30% when the pitch is doubled. Thus, a 3-mm scan with a pitch of 2 has an effective slice thickness of 3.9 mm. Apply this concept to the following situation:

What technical parameters provide the best scan quality if the distance to be scanned is 10 cm? If one uses 5-mm collimation with a pitch of 1, the scan will take about 20 seconds (depending on tube rotation time) with an effective slice thickness of about 5 mm. If 3-mm collimation with a pitch of 2 is employed, the scan will require just under 17 seconds with a 3.9 mm effective slice thickness. Thus, whenever possible, it is preferable to maximize spatial resolution using thinner collimation while increasing the pitch to cover the entire scan distance more expediently. a faster scan will increase scanner throughput and minimize respiratory motion degradation.

Field-of-View

The field-of-view (FOV) deserves special consideration for two reasons. First, because the conditions that prompt helical CT examination of the aorta are often nonspecific, evaluation of the entire thorax is important. the diagnosis may be found in the pleural space or peripheral lung parenchyma, not the aorta. Additionally, using small FOVs will decrease photon flux in the imaging volume, contributing increased image noise. This may be partially compensated for by increasing the mA, but the scanner may not permit a sufficient increase in ma if the length of the scan is long. for these reasons, it is advisable to measure the FOV from outer rib to outer rib at the widest portion of the thorax. This FOV creates a good compromise between a large enough FOV to visualize the entire thorax yet provides a larger image with improved resolution that is easier to read.

Reconstruction Increment

A reconstruction increment that provides nearly a 50% overlap in slice thickness is generally sufficient to generate excellent quality images. Recent work suggests that there is limit at which overlapping the reconstructions no longer provide improved longitudinal resolution.^[1] Beyond this point, further overlap only generates more images and thus only increases scan reconstruction time and storage requirements. the optimal overlap depends on the pitch employed. As a compromise, a 50% overlap generally provides optimal resolution and enhances the quality of postprocessing techniques. One should avoid using reconstruction increments that are exact harmonics of the collimation employed (i.e., a 2.5-mm reconstruction increment with a 5-mm collimation scan). in such circumstances, artifacts may be introduced if post-processing techniques are used.

Iodinated Contrast

Amount and Concentration: There is no consensus regarding what iodine concentration should be routinely employed for helical CTa of the aorta. the common practice of using undiluted (300 mg I/mL or 360 mg I/mL) is usually sufficient. in one study, an iodine concentration of 125 mg I/mL was found to be unacceptable at any flow rate.^[2] in this same study, an iodine concentration of 150 mg I/mL provided satisfactory images, when the flow rate was adjusted. This iodine concentration had the added benefit of decreasing the amount of streak artifact emanating from thoracic venous structures. However, questions regarding the safety of diluted contrast preparations remain because such dilute mixtures are not generally commercially available. Obviously, there are issues of practicality as well. Generally, we employ undiluted contrast (300 mg I /mL) for all thoracic helical CTa applications.

Rate: There is considerable variation among investigators regarding the rate at which contrast is delivered for helical CTa of the thorax. Generally speaking, a rate of 3 cc/sec is sufficient, although some investigators use rates as high as 5 cc/sec. Some consider higher rates as less practical because they may require larger bore catheters; however, rates of 4 to 5 cc/sec may be achieved using 20-gauge catheters. Generally, an injection rate of 3.5 cc/sec is satisfactory for helical CTa applications in the chest. for such injection rates, 22-gauge catheters are adequate. Indeed, flow rates of 4 to 4.5 cc/sec may be used through such catheters when required. It should be noted that when higher injection rates are used, a larger volume of contrast must be delivered to maintain the injection throughout the scan acquisition.

Injection Timing: An appropriate scan delay is critical for excellent vascular imaging. As a general rule of thumb, a standard delay of 25 seconds for helical CTa of the thoracic aorta will usually suffice. Patients with cardiomyopathy or an unsuspected stenosis of the injected vein may require longer delays to achieve optimal aortic opacification. Patients with hyperdynamic cardiac function may require a shorter delay. the time to peak aortic opacification from patient to patient varies considerably. Van Hoe et al^[3] found that this delay varied from 11 to 30 seconds, with an average of 20 seconds. One may perform a test injection to determine the time to optimal aortic opacification in a given patient. With this method, a 20-cc bolus of contrast is administered, followed by a 10 second delay. After the 10 second delay, one image every 2 seconds is acquired at the same level (often the cranial aspect of the descending thoracic aorta) for a total of 30 seconds. the image with the greatest contrast density is used to select the proper scan delay. This method is somewhat cumbersome to perform routinely, and has the added detraction of an increased radiation dose as well as a slight increase in visceral background attenuation due to the injected contrast.

As an alternative, bolus timing software packages may be purchased for most modern CT scanners. Such programs allow the time to peak enhancement to be monitored without a separate test injection. These programs monitor the attenuation of the vessel in question during the contrast injection, and display the attenuation graphically in real-time. Once the graph demonstrates a sharp rise in attenuation, the scan sequence is triggered manually. When possible, we prefer to use bolus timing software to ensure proper arterial opacification. With even minimal experience, the CT technologist can learn to achieve excellent arterial opacification without the direct super-vision of the radiologist.

Distance to Be Scanned

The entire distance to be scanned is important for several reasons. Perhaps, most importantly, the contrast injection must be maintained for all but the last few seconds of the scan acquisition. Failure to do so may allow unopacified blood to enter the imaging volume, thus creating artifacts and diminishing the quality of the study. Second, the longer the distance to be scanned, the greater the heat loading on the tube, necessitating a lower ma be used. This is particularly problematic in larger patients, who may require all the photons the tube can muster. Third, imaging beyond the desired structures will provide additional information, but at the expense of more tube exposures. As the x-ray tube has a finite life span, extra tube exposures may come at significant cost. Finally, imaging beyond the desired volume also contributes to increased patient radiation dose.

kVp and mA

There is usually no reason to adjust the kVp from the standard values of 120 to 140 kVp. However, the ma used is critically important for achieving excellent quality studies. While it is preferable to keep ma as low as possible to reduce patient dose, little is accomplished if the patient receives a somewhat lower radiation dose at the expense of a poor quality study. the contrast resolution on CT scanners is heavily dependent on mA. Noisy images resulting from increased quantum mottle often result from using ma values that are too low for a particular patient. Although for the average patient, ma is often not a major consideration, for larger patients, an upward adjustment in the ma is critical to achieve a high quality study. Often this adjustment must be performed manually because modern CT scanners are often programmed to minimize the ma to reduce tube heat loading and patient dose. When larger patients are being scanned, the ma should be adjusted manually to near maximum to diminish image noise and provide the highest quality imaging. Careful attention to FOV and scanning distance in this circumstance is also required.

Tube Rotation Time

Modern scan-ners now allow tube rotation times as low as 0.5 seconds. GE scanners (GE Medical Systems, Milwaukee, WI) have a minimum tube rotation time of 0.8 seconds. There are options to increase the tube rotation time that may have a role in other thoracic applications (e.g., HRCT). Helical CTa of the aorta, however, is best accomplished with the fastest allowable tube rotation time.