Computed Tomographic Angiography

Major Components of a Computed Tomography Scanner

The major components of a CT scanner are an x-ray tube and generator, a collimator, and photon detectors. These components are mounted on a rotating gantry which produces the x-rays necessary for imaging. The predetector collimator helps shape the beams that emanate from the x-ray tube in order to cut out unnecessary radiation. The detectors consist of multiple rows of detector elements (> 900 elements per row in the current scanners), which receive x-ray photons that have traversed through the patient, with the postdetector collimators preventing backscatter which degrades image quality. The newer scanners have as many as 320 detector rows and the width of each detector (“detector collimation”) has decreased from 2.5 mm in 4-slice systems to 0.5 mm. The most important benefit of increasing the detector rows is the increased coverage per gantry rotation (a 320-row detector CT with a detector width of 0.5 mm will have 160 mm z-axis coverage). The submillimeter detector width improves spatial resolution in the z-axis while increased coverage shortens the scan time. Each detector element contains radiation-sensitive solid state material (such as cadmium tungstate, gadolinium oxide or gadolinium oxysulfide), which converts the absorbed x-rays into visible light. The light is then detected by a silicone photodiode, amplified, and converted into a digital signal. The gantry rotation time determines the temporal resolution of the images with older scanners having a rotation time of 0.75 second while the more contemporary scanners have a rotation time of 0.33 second. The temporal resolution of a single-source scanner, one x-ray generator mounted on the gantry, is slightly higher than half the time it takes for the gantry to rotate 360 degrees. Thus a 0.33 second gantry rotation will effectively provide a temporal resolution of 0.17 second. With a dual-source scanner, two x-ray generators mounted on the gantry, the temporal resolution will improve by a factor of 2.

Computed Tomography Attenuation Data for Image Reconstruction

CT measures the local x-ray attenuation coefficients of the tissue volume elements, or voxels, in an axial slice of the patient’s anatomy. The attenuation coefficients are then translated into gray-scale values (CT value) of the corresponding picture elements (pixels) in the displayed two-dimensional image of the slice. The numerical CT value is normalized to the attenuation properties of water and is reported as “Hounsfield units” (HU). Pixel values are stored as integers, in the range–1024 HU to 3071 HU, corresponding to 4096 different gray-scale values. By convention water = 0 HU, air = −1000 HU, and is independent of the x-ray spectrum. The CT values of human tissue, however, depend on the x-ray spectrum. In general, lung and fat have negative CT values, muscle has a positive HU and bone has rather large CT values up to 2000 HU. Administration of iodine contrast agent increases the CT value with contrast-filled vessels typically having CT values in the range 200 HU to 600 HU. In most cases contrast-filled vessels can be easily differentiated from the surrounding tissue, which does not exceed a CT value of 100 HU with the exception of bone. This easy, threshold-based differentiation is the basis for CTA and related image postprocessing techniques. The gantry rotates around the patient collecting attenuation data from different angles. The attenuation coefficient also varies depending on the energy of the photons (measured in keV). The measured intensity of photons at the CT detector is related to the photon flux (number of photons) coming through the x-ray tube and that detected at the detector elements. Consequently, as the attenuation of the tissue increases, the fraction of photons that are detected at the detector element decreases. Photon energy (keV) and photon flux (milliamperes [mA]) are variables that are set by the user. Increasing tube current (mA) will improve image quality at the expense of increasing radiation dose. Certain manufacturers have introduced an “effective” mAs concept for spiral/helical scanning, which incorporates the amount of time the tube current is being generated. Tube voltage (kV) determines the energy of the x-ray beam or the hardness of the x-ray. A higher kV results in a smaller fraction of the x-ray beam being absorbed (reduced attenuation) but will result in improvements in contrast.

Scanning Modes

The two scanning modes used in CT are the axial mode and the spiral/helical mode. The major differences between these modes include (1) differences in table movement during image acquisition, (2) differences in assignment of data to each channel, and (3) need for interpolation for data reconstruction. Each mode has its benefits; however, the mode used for vascular CTA is the spiral/helical mode. For coronary CTA, there has been a shift toward using the axial mode due to its benefit in significantly reducing radiation exposure.

Beam Pitch

Pitch is an expression of the relationship between the table distance moved per gantry rotation and the coverage of the scanner. Pitch = (table feed per gantry rotation [mm]/coverage [mm]). If the pitch is 1 then there would be no gaps between the data set; however, if the pitch is greater than 1, gaps would be present, and if the pitch is less than 1 there would be overlap in the data acquisition. The pitch for electrocardiographic (ECG) gated cardiac scanning is typically 0.2–0.3, whereas for vascular CT the pitch ranges between 0.5–1.2.

General Acquisition Parameters

The selection of the specific acquisition parameters of imaging depends on the employed scanner model, the patient’s body habitus, and the clinical question. The two main adjustable parameters are the tube voltage and current. The voltage is typically set at 120 kV although 100 kV provides acceptable images with significantly reduced radiation and can be employed in most individuals for vascular imaging who are not obese. Tube current is usually 200–300 mA and again can be adjusted upward if the patient is very large. Breath-holding is required for chest and abdomen CTA acquisitions in order to reduce motion artifact. In MDCT spiral scans, the volume coverage speed ( v , cm/s) can be estimated by the following formula:

where M = number of simultaneous acquired slices, s _coll = collimated slice-width, p = pitch, and t _rot = gantry rotation time. Although current generation scanners offer improved spatial resolution, their increased coverage and rotation speeds pose the risk of “out-running” the bolus of contrast in CTA applications. Accordingly, adjustments in both the pitch and the gantry rotation speed must be made to achieve a table translation speed of no more than 30–32 mm/s for CTA applications. In a 64-slice scanner, this usually is achieved by a reduction in t _rot to 0.5 second and a decrease in pitch to ≤ 0.8.

Electrocardiogram Gating

ECG gating is a method of gating imaging events to portions in the cardiac cycle where motion may be minimal, namely diastole. ECG gating is indispensable for coronary imaging and vascular structures that are prone to cardiac motion artifact such as the ascending aorta. The two most common ECG gating methods are retrospective and prospective gating. With traditional spiral scanning the ECG gating is performed retrospectively where the data and ECG information are acquired and subsequent reconstructions can be performed at various time points in the R-R interval. Compared to prospective ECG gating, which is the method used for axial scanning, the scan is triggered at the R wave and image acquisition occurs at a fixed point in the cardiac cycle. However with recent advances in CT imaging, it is also now possible to perform a prospective ECG triggered helical scan using high pitch with extremely low radiation exposure. These have been referred to as “flash” scans and are gaining significant popularity for coronary imaging.

Contrast Administration

All angiographic x-ray contrast remains in the extracellular space and rapidly distributes between the intravascular and extravascular spaces immediately after intravenous administration. It is the process during the early phase of rapid contrast distribution and redistribution that determines the vascular enhancement. Vascular enhancement differs significantly from parenchymal (soft tissue) enhancement characteristics. The two key components that determine arterial enhancement are the amount of contrast per unit time (mL/s) and the duration of administration (seconds). The resulting product of the two is the volume of contrast (flow rate × duration). For example, 100 mL of contrast media given at 5 mL/s will require 20 seconds to deliver. The relationship between flow rate, volume of contrast, and duration of administration is the most important concept to understanding injection protocols for vascular imaging.

Currently, low- or iso-osmolar nonionic contrast agents are the most commonly used for CTA. It is imperative to assess renal function prior to administration of contrast so decisions can be made in regard to prophylactic measures, type of contrast used, and whether the study should be cancelled. The contrast is given intravenously using a power injector. Since contrast arrival time to the region of interest may vary, appropriate timing needs to be determined by using a test bolus or automated bolus tracking technique. The less commonly used technique of using a test bolus is performed by giving a small dose of contrast material and determining the time it takes for the region of interest to opacify. More commonly, a triggered or automated bolus tracking technique is used where a region of interest is drawn on the aorta closest to the area of interest. A repetitive low-dose acquisition is acquired 5 to 10 seconds after contrast administration until an HU threshold is achieved (typically 110 HU). The actual CTA will be acquired once this threshold is obtained. The typical volume of contrast used is 100–120 mL with an iodine concentration between 320–370 mg/mL administered at a rate of 4 mL/s followed by a saline flush.

Image Reconstruction at the Scanner Console

There are various image reconstruction filters offered by each manufacturer. Filters are referred to as “sharp” or “soft” filters. Sharper reconstruction filters will provide more details but also more noise and are best for assessment of stents and areas of calcification. Softer reconstruction filters provide less image detail but less noise as well. Soft to medium filters are usually used for most CTA applications. Image reconstruction can also be performed at different cardiac phases of cardiac gated acquisitions. It may be important in assessment of coronary anatomy in cases of thoracic aortic dissection and thoracic aortic aneurysms. This is most important for cardiac CTA where coronary anatomy may need to be assessed at different phases to ensure accurate delineation of coronary stenosis. When ECG gated thoracic aortic imaging is performed, various phases can be reconstructed to assess the aorta.

Slice width and slice increment used for image reconstruction at the scanner console depends on the anatomy being assessed and scanner capabilities. Reconstruction thickness for vascular imaging can be performed at the same width (thin) or several times the detector width (thick) to reduce noise. Thinner slices are associated with higher image noise compared to thicker slices and take longer time to review. A slice increment of approximately 50% of the slice thickness is typically used.

Image Postprocessing

Similar to vascular magnetic resonance angiography (MRA) (see Chapter 13 ), multiple postprocessing techniques can be used in vascular CTA to assess the hundreds to thousands of images that are generated. Usually two data sets are reconstructed including “thick” and “thin” sets. The thick set (5.0 mm) is used for general assessment, whereas the thin set (0.5–0.75 mm) is better suited for detailed evaluation. Image formats used for evaluation include (1) multiplanar reformats (MPR), (2) maximal intensity projections (MIP), (3) curved planer reformats (CPR), (4) volume rendering (VR), and (5) shaded surface display (SSD). (See Chapter 13 for a description of these techniques.) For CTA, the evaluation of the data set begins with review of the axial images to assess gross anatomy and scan quality. A MIP format is used to view the vascular structure of interest in the traditional projections as well as in oblique orientations. A major caveat is the presence of calcium when viewing MIP images, as it can overestimate the severity of stenosis. For detailed evaluation, especially when calcium and/or stents are present, the raw MPR images need to be reviewed. Curved planar reconstruction is a unique technique that makes it possible to follow the course of any single vessel and displays it in a nontraditional plane where the entire vessel can be seen in a single image. 3D-VR images can also be reviewed to get a general appreciation of the anatomic variations if necessary. Each of these reconstruction methods has its pitfalls and it is important to develop a systematic process to identify and evaluate an abnormality.

Computed Tomography Specific Dosimetry

In addition to the nomenclature for radiation dosimetry described above, a particular set of terms has been developed for CT. The dose profile (D[z]) for a CT scanner is a mathematical description of the dose as a function of position on the z axis (perpendicular to the tomographic plane). The CT dose index (CTDI), measured in units of Gy, is the area under the radiation dose profile for a single rotation and fixed table position along the axial direction of the scanner, divided by the total nominal scan width or beam collimation. CTDI is difficult to measure and therefore not commonly reported. Instead, the CTDI 100 is measured. CTDI 100 represents the integrated radiation exposure from acquiring a single scan over a length of 100 mm. To estimate the average radiation dose to a cross-section of a patient’s body, a weighted CTDI (CTDI _w ) is calculated. This is determined by the equation:

CTDI _w , given in mGy, is always measured in an axial scan mode and depends on scanner geometry, slice collimation, and beam pre-filtration as well as x-ray tube voltage, tube current (mA), and gantry rotation time ( t _rot ). The product of mA and t _rot is the mAs-value of the scan. To obtain a parameter characteristic for the scanner used, it is helpful to eliminate the mAs-dependence and to introduce a normalized ( CTDI _w ) _n given in mGy per mAs:

The important CT-specific dosimetry term is the volume weighted CTDI , or CTDI _vol . This quantity represents the average radiation dose over the volume scanned in a helical or sequential sequence. It is determined from the CTDI _w by the equation:

CTDI _vol is used to determine the dose-length product (DLP) , measured in units of mGy ∙ cm. DLP reflects the integrated radiation dose for a complete CT examination, and is calculated by:

Many CT scanner consoles report the CTDI _vol and DLP for a study. DLP can be related to E by the formula:

where E _DLP , measured in units of mSv/(mGy ∙ cm), is a body region-specific conversion factor. The most commonly used E _DLP values are reported in the 2004 CT Quality Criteria. These E _DLP values are determined using Monte Carlo methods, averaged for multiple scanners.

Dose Reduction Techniques

There are several ways to lower the dose delivered to a patient. These methods can be used either in isolation or combined to lower the exposure exponentially. Reducing tube current will lead to a direct reduction in the radiation dose to a patient. However, a conscious decision needs to be made on whether the trade off on radiation reduction outweighs image quality. This becomes very important in obese patients where the reduction in tube current may result in rather poor images. The contrast to noise ratio increases with decreasing x-ray tube voltage. As a consequence, to obtain the adequate contrast to noise ratio, the dose to the patient may be reduced by lowering the kV setting. There is nearly a 50% reduction in the radiation exposure when using 80 kV instead of 120 kV when performing CTA. A recent study recommends 100 kV as the standard mode for aortoiliac CTA and reports dose savings of 30% without loss of diagnostic information.

ECG-controlled dose modulation is a method that is employed during continuous imaging with retrospective studies. Typically, the output is kept at its nominal value during a user-defined phase (in general the mid- to end-diastolic phase) while during the rest of the cardiac cycle, the tube output is reduced to 20% of its nominal value to allow for image reconstruction throughout the entire cardiac cycle. Using this technique, dose reduction of 30% to 50% has been demonstrated in clinical studies.

Anatomical tube current modulation is a technique adapted to the patient geometry during each rotation of the gantry. The tube output is modulated on the basis of the tissue attenuation characteristics of the localizer scan or determined online by evaluating the signal of a detector row. By employing such a technique, dose can be reduced by 15% to 35% without degrading image quality depending on the body region. A more sophisticated variation of anatomic tube current modulation varies the tube output according to the patient geometry in the longitudinal direction in order to maintain adequate dose when moving to different body regions, for instance from thorax to abdomen (automatic exposure control). Automatic adaptation of the tube current to patient size prevents both over- and under-irradiation, considerably simplifies the clinical workflow for the technician, and eliminates the need to look up tables of patient weight and size for adjusting the mAs settings.

Dual Energy Computed Tomography

Within the last few years, certain technological advancements have occurred including dual energy and spectral CT. Standard CT contains a polychromatic range of x-ray photon energies. Lower level photons exhibit the photoelectric effect while higher energy levels produce the Compton effect. The photoelectric process predominates when lower energy photons interact with materials of high atomic number and Compton scattering predominates at higher photon energies with materials of lower atomic numbers.

Spectral CT or dual energy CT takes advantage of different x-ray photon energy levels to highlight differences in tissue properties. Several photon energies can be used, but typically a low and high energy profile are utilized to highlight differences in tissue attenuation (80 kVp vs. 140 kVp and hence “dual energy”). Siemens provides a dual source-detection scanner that operates at two different energy levels. General Electric provides a single source detector that switches between high and low tube potentials, while Phillips provides a single x-ray source that operates at a constant polychromatic range and that uses a double layer of detectors to produce both low and high energy images.