Computed Tomography

Computed Tomography


Computed tomography (CT) has experienced enormous growth in clinical use over the past three decades primarily due to significant advances in image quality and a dramatic reduction in acquisition time as the technology has advanced. Image quality has increased because of better detector sampling along the long axis (z-dimension) of the patient with multiple detector array systems of 40 to 160 mm beam coverage in one rotation. CT reconstruction has advanced from filtered backprojection (FBP) to generations of iterative reconstruction and deep learning (DL)-based adjuncts, providing lower noise images at lower radiation dose levels for improved spatial resolution and contrast resolution. With half-second gantry rotation (or shorter), 80 mm of patient length can be acquired in 1 second (s), and hence 400 mm of patient length can be scanned in 5 s.

CT imaging has gained widespread use across many clinical applications for abdominal, thoracic, head, musculoskeletal, CT angiography (CTA), organ-specific CT perfusion, and dual-energy CT. Thin-slice axial images, multiplanar reconstruction, and 3D volume rendering reconstruction allows segmentation of bone structures from soft tissue, accurate assessment of iodinated contrast in CT angiograms, and the ability to use volumetric data for 3D printing. Coronal and sagittal CT visualization provides important additional information to the interpreting physician, for spinal alignment, orientation of pathology, and normal anatomy including identifying feeding vessels, gastrointestinal tract topography, abdominal organ placement and orientation, trauma, and other factors.

  • 1. Acquisition geometry—x-ray tube rotates in a gantry around the long axis of the body

    • a. For one rotation, multiple projections are acquired and reconstructed into axial images.

    • b. Two-dimensional images are comprised of pixels with equal dimensions in horizontal (x) and vertical (y) axes.

    • c. The image represents a volumetric section of the patient with a thickness Δz (Fig. 10-1).

    • d. Dimensions of each voxel are on the order of 0.6 mm (Δx and Δy) by 0.5 mm (Δz).

    • e. A 75-kg person can be imaged with over 400 million voxels using CT.

  • 2. Volume data set and multiplanar reconstruction from thin-slice axial images (Fig. 10-2)

    • a. Routine reformatting into the coronal and sagittal planes provide intuitive anatomical display.

FIGURE 10-1 Individual CT axial images are comprised of a two-dimensional matrix of pixels with dimensions of Δx and Δy. Each pixel in the image corresponds to a voxel in the patient. The Δz dimension can be of similar or different size to the Δx or Δy depending on acquisition parameters. image F10-3

FIGURE 10-2 Multiplanar reformatting of the axial voxels into the coronal and sagittal planes provides radiologists with views of continuous anatomical features with resolution similar to the inplane axial images because of the ability to acquire slices with a small Δz dimension. image F10-4


  • 1. X-ray tube potential for routine scanning—120 kV; other tube voltages also used

    • a. Tube voltage is varied to optimize image quality versus radiation dose for clinical applications.

    • b. Protocols and selection of voltage are based on patient size and clinical need or diagnosis.

    • c. Tube voltage options (typical): 80, 100, 120, 140 kV (Fig. 10-3).

      • (i) Added filtration in beam creates a “hard” x-ray spectrum.

      • (ii) Physical properties (attenuation) of tissues are based on these spectra.

    • d. Effective energy for kV spectra range from approximately 43 to approximately 70 keV (Fig. 10-4).

    • e. Compton scatter is 10-fold more likely than photoelectric or Rayleigh scatter interactions for soft tissue.

    • f. For bone and iodinated contrast agents, the photoelectric interaction does play an important role.

  • 2. Attenuation coefficient for Compton interactions, Equation 10-1:

    where N is Avogadro’s number (6.023 × 1023), Z is the atomic number, and A is the atomic mass

    FIGURE 10-3 The x-ray spectra for computed tomography are shown for 80, 100, 120, and 140 kV—each spectrum is filtered with 10 mm of aluminum, which approximates the typical amount of filtration at the center of the beam for most CT scanner models. The high tube potentials along with the large filtration thicknesses lead to x-ray beams with high effective energy. image F10-5

    FIGURE 10-4 What physical parameter does the grayscale in a CT image correspond to? The mass attenuation coefficient of soft tissue is shown as a function of x-ray energy in this graph. The curves correspond to the attenuation coefficients for the photoelectric effect, Rayleigh, and Compton scattering. The effective energy in CT is outlined by the vertical band. In this region, it is seen that the Compton cross-section is approximately 10-fold that of the Rayleigh or photoelectric cross-sections, meaning that the grayscale (Hounsfield unit) in CT for soft tissue is primarily determined by the physical dependency of Compton scattering—electron density. image F10-6

    • a. For most elements (except H), Z/A ratio is ½; density tends to dominate when comparing adipose-rich tissues (lower density) to soft tissues in terms of attenuation at these energies.

  • 3. The Hounsfield unit (so named after a principal developer of CT, Sir Godfrey Hounsfield)

    • a. Preprocessing of CT acquired data applies a logarithmic function to the ratio of incident to transmitted x-ray intensity: ln image; this effectively linearizes the effective attenuation coefficient.

    • b. Subsequent reconstruction provides grayscale encoding (Eq. 10-2), a function of the linear attenuation coefficient, called the Hounsfield unit (HU).

      For a given voxel in the image, K, which contains average µK, the HU is scaled to µw, the linear attenuation coefficient of water.

    • c. The HU is defined at water (µK = µw); HUwater = 0, and air (µK = µair); HUair = -1,000.

    • d. µ has relatively strong dependencies on x-ray beam energy as µ(E).

      • (i) The notation for µ in Equation 10-2 represents the effective linear attenuation coefficient, µeff.

      • (ii) Calibration of HU is slightly different at each tube voltage, but for water HU = 0 at all tube voltages; other tissues such as liver and bone will have slightly different HU values at different tube potentials.

    • e. Imprecision in HU values results from calibration, scattered radiation, quantum noise, and beam hardening.

    • f. Typical HU for water is usually within ±5 HU with often larger deviations for other tissues and air.


  • 1. The gantry—geometry and detector configuration

    • a. Clinical CT scanners have the x-ray source and detector arrays arranged in an arc (Fig. 10-5).

    • b. Detector arrays aligned along a radius of curvature provide perpendicular incidence of the beam.

    • c. A fan-beam projection incident on individual detectors.

      • (i) For a given view (tube position and view angle), the individual x-rays (rays) define a line integral extending from the source to each individual detector along the detector array.

      • (ii) The collection of all rays across the detector elements constitutes a view or projection.

      • (iii) Raw data collected during a scan are comprised of rays and views.

    • d. Basic geometry of a CT scanner and dimensions (Fig. 10-6).

      • (i) Isocenter of rotation is almost always the center of the reconstructed image.

        FIGURE 10-5 The general configuration of a CT scanner is illustrated in cross-section. The x-ray tube is mounted on a rotating gantry, onto which the x-ray detector arrays are also mounted, in the typical rotate-rotate (third-generation) geometry in a fixed geometrical orientation. A beam of x-rays is projected through the patient to the detector and the raw acquired signal is recorded. The line between the x-ray source and each individual detector element (dexel) corresponds to the fundamental measurement in CT, a ray. The group of rays corresponding to one fan-beam projection is called a view. image F10-7

        FIGURE 10-6 The basic geometrical dimensions of a modern rotate-rotate CT scanner are shown. The detector arrays are located on a radius-of-curvature positioned at a distance B from the x-ray source. The source-to-isocenter distance is given by a distance A, and hence the magnification factor (M) of objects at iso-center is given by the ratio M = B/A. For most modern CT scanners, the fan angle is about 50° to 60°, and combined with the other geometric parameters this defines the maximum field-of-view (FOV). image F10-8

      • (ii) For source to isocenter distance A, and source to detector distance B, magnification M = B/A.

      • (iii) Detector width and length specified at isocenter: if B = 95 cm and A = 50 cm, M = 1.9 and detector width specified as 0.5 mm has an actual physical width of 1.9 × 0.5 = 0.95 mm.

      • (iv) Fan angle of the x-ray beam is approximately 50° to 60° and defines maximum scan FOV.

    • e. Multidetector array CT scanner (MDCT) geometry (Fig. 10-7).

      • (i) 16, 64, 256, or more detector arrays are abutted to create a larger sampling along the z-axis.

      • (ii) Divergence of the x-ray beam in the z-direction results in a cone angle.

      • (iii) Example: a 64 detector array of 0.625 mm width measures 40 mm along z, requiring a similar beam width obtained by opening the collimator slit at the x-ray tube port.

      • (iv) Systems allow detectors to bin data from adjacent detector arrays to increase slice thickness.

      • (v) Larger slice thickness: less image noise (both electronic and quantum), less storage required.

      • (vi) Larger slice thickness: more partial volume artifact, less z-axis resolution.

      • (vii) MDCT reconstructed slice thickness and collimated x-ray beam width are separate parameters.

      • (viii) MDCT allows acquisition of faster scans (whole body in seconds for large width arrays). (ix) MDCT allows thinner slice acquisitions (thinnest slices are sometimes termed “raw” data).

    FIGURE 10-7 The angles associated with a modern rotaterotate CT scanner using multidetector arrays are shown. The fan angle is defined between the x-ray source and the extent of the detector arrays in-plane. All modern multiple detector array CT scanners (MDCT) can be considered cone-beam CT systems, with full-cone angles of 4.2° for a 40-mm detector width at isocenter, and 17.2° for a 160-mm detector width. This cone angle is fully considered in the reconstruction algorithm used to convert the raw acquired data into the final CT volume data set. image F10-9

  • 2. Wide-beam (cone-beam) CT systems

    • a. All whole-body scanners with ≥64 detector channels are technically cone-beam CT systems.

    • b. True cone-beam scanners: where the half cone angle approaches 9°.

    • c. Reconstruction algorithms for synthesizing data must take into account the cone angle.

    • d. Challenges and benefits of wide cone-angle CT systems.

      • (i) Challenges: increased x-ray scatter detection and cone-beam artifacts

      • (ii) Benefits: whole organ imaging without table motion (e.g., head, heart, organ perfusion; the latter with high temporal resolution)

    • e. Cone-beam systems with flat-panel detectors (Fig. 10-8).

      • (i) Full cone angles almost as large as the fan angle are possible.

        FIGURE 10-8 While modern whole-body CT systems are rightly considered as (small cone angle) cone-beam systems, more traditional specialty application cone-beam CT systems typically use a flat-panel detector opposed to an x-ray source in a rotate-rotate (third-generation) cone-beam geometry, as illustrated. The system in this diagram is designed for pendant-geometry breast imaging and rotates in the horizontal plane. Flat-panel-based cone-beam CT systems enable dedicated CT applications in orthopedic, dental, breast, radiation therapy positioning, and angiographic applications and typically provide higher spatial resolution images albeit with greater artifact potential. image F10-10

      • (ii) Full 2D bank of detectors and flat geometry, with different reconstruction geometry.

      • (iii) Detector calibration includes flat-fielding to correct for inverse square variations and heel effect.

      • (iv) Parallax error (nonnormal incidence to the detector at wide angles) cannot be corrected.

      • (v) Detector readout rates limit acquisition speed; pulsed x-ray tube projections limit motion.

      • (vi) Applications: dental CT, breast CT, orthopedic, angiographic, and radiation therapy.

  • 3. The slip ring

    • a. The CT gantry x-ray tube and detector array rotate continuously in one direction with a slip ring.

    • b. The slip ring has a number of electrically conductive tracks and associated contactors (Fig. 10-9).

      • (i) This maintains electrical contact with the stationary frame to the rotating gantry.

      • (ii) The x-ray transformer is mounted on the gantry to avoid high voltage and electrical arcing.

      • (iii) Numerous small conductive tracks with low electronic noise transfer digital signal data out.

      FIGURE 10-9 A CT scanner involves a rotating frame (the gantry), as well as a stationary frame that includes mechanical and electrical components. To convey power onto the rotating gantry from the stationary frame, as well as to conduct signal data from the rotating gantry to the stationary frame, a slip ring is used. A slip ring uses gliding contacts to allow communication and power transfer between the stationary and rotating frames without the use of wires, and this enables the gantry to rotate continuously in a single direction. Previous generations of CT systems did not use slip rings and were bound by mechanical cable-based connections, substantially limiting gantry rotation rates. Slip rings have enabled gantry rotation periods to move from 3.0s (when cables were used) to modern CT rotation periods of as little as 0.25s. The power transfer components of the slip ring may involve the conductance of approximately 50 A at approximately 400 V—so accidental contact would be deadly. image F10-11

    • c. Helical and high-speed rotations are enabled, with high centrifugal forces exceeding 20 g’s.

      • (i) Special design considerations and hardware are required for all components on the gantry.

      • (ii) X-ray tube focal spots have flat emitter design with magnetic/electrostatic focusing coils to withstand high gyroscopic forces (see Chapter 6 for details).

  • 4. The patient table (couch)

    • a. The table is an important and highly integrated CT scanner component.

    • b. The CT computer controls longitudinal table motion with precision motors and feedback.

      • (i) Precise control and positioning of the table is important, particularly for helical acquisition.

      • (ii) Accuracy of anatomic positioning is essential.

      • (iii) Table position accuracy is evaluated during routine CT scanner testing.

    • c. Height adjustment allows for convenient patient access and centering in the FOV (Figs. 10-10 and 10-11).

    FIGURE 10-10 The patient table is a surprisingly high-tech component of a CT scanner. The patient table lowers to sitting height to allow patients—including the elderly and physically impaired—to sit on the table and reposition to a prone or supine position, with help from the attending technologist. image F10-12

    FIGURE 10-11 Once the table is raised to the appropriate height (to isocenter), the table becomes the essential linkage between the translation of the patient along the z-dimension during acquisition and the position of the resulting CT image data set. Hence, the translational accuracy of the table is an essential component of quality control procedures. image F10-13

  • 5. The x-ray tube and CT detector characteristics

    • a. CT x-ray tubes have the highest power loading and heat dissipation ratings.

      • (i) MDCT acquisition geometry: better use of tube output with larger collimator and beam width.

      • (ii) Plane of anode rotation is parallel to plane of gantry rotation to reduce gyroscopic forces.

      • (iii) Anode-cathode axis is parallel to the z-axis of the scanner—so is the heel effect (Fig. 10-12).

      • (iv) The x-ray output is limited angularly in the anode-cathode direction, but not in the fan angle (Fig. 10-13).

      FIGURE 10-12 A. The plane of the x-ray tube anode is parallel with the rotation of the x-ray tube around the gantry, as illustrated. The high rotational velocity and the large mass of the x-ray tube anode creates large gyroscopic forces, and when considered with the high rotational velocity of the gantry, the only practical mechanical orientation of the x-ray tube is to have the rotational planes of the anode and the gantry be parallel. B. With the orientation of the x-ray tube and the gantry as defined in (A), the anode-cathode direction is aligned with the z-axis of the CT scanner. Hence, the heel effect (always parallel to the anodecathode axis) runs along the z-dimension of the scanner, which is also the dimension of table/patient travel corresponding to the vertical dimension of coronal and sagittal CT images. image F10-15

      FIGURE 10-13 The x-ray tube anode and the orthogonal dimensions of CT acquisition are displayed. A. The x-ray beam plane emanating from the plane of the anode constitutes the fan beam, laterally across the fan angle. B. The beam thickness dimension extends across the anode-cathode plane and corresponds to the cone angle of the scanner. image F10-16

      FIGURE 10-14 The rotating CT gantry (A) produces motion blur (rotational velocity of +v). B. To partially address this, some CT systems steer the focal spot (by steering the electron beam between the cathode and anode) in the opposite direction (-v) of gantry motion (C). This is accomplished with magnetic deflection of the electron beam incident on the anode. image F10-17

    • b. Continuous x-ray output is used by most CT scanner manufacturers.

      • (i) Detector sampling time is the acquisition time for each CT projection acquired.

      • (ii) Sampling dwell times: approximately 0.2 to 0.5 ms.

      • (iii) Approximately 1,000 to 3,000 projections are acquired for a 0.5 s gantry rotation over 360°.

      • (iv) At the periphery of the FOV, the angular displacement of the focal spot during the sampling dwell time is greater than at the center, causing a relatively greater sampling distance, and resulting in blurring and loss of spatial resolution.

      • (v) Compensation (some CT scanners) steers the focal spot in the opposite direction (Fig. 10-14).

      • (vi) Some CT scanner manufacturers use focal spot steering to “oversample” the z-dimension to cause a shift in the source position—this effectively doubles the unique angular projection data sampling per rotation, relative to the number of detector channels.

    • c. Pulsed x-ray output by some manufacturers is used to switch x-ray voltage for dual-energy CT.

      • (i) Every other pulse is a high (e.g., 140 kV) or low (e.g., 80 kV) projection (see Section 10.3.8).

    • d. Wider x-ray beam collimation with MDCT increases the detected scatter fraction.

      • (i) 2D antiscatter grids are used to reduce scatter (Fig. 10-15).

    • e. Indirect detection (scintillator) solid-state detectors are most widely used (Fig. 10-16).

      • (i) Scintillator materials are sintered to increase density and improve detection efficiency.

      • (ii) Ceramic phosphors are scored to produce individual detector elements.

      • (iii) Opaque filler between detector elements is added to reduce cross-talk.

        FIGURE 10-15 Scattered radiation striking the x-ray detector has substantially increased with wide-beam collimation. A two-dimensional high grid ratio system is employed. Primary x-ray attenuation by the antiscatter grid is reduced by aligning the septa with the dead spaces in the detector array. image F10-18

        FIGURE 10-16 A. A cross-section of a ceramic wafer (e.g., Gd2O2S) is layered onto a photodiode. Slits between individual detector elements are filled with an opaque material to eliminate optical cross-talk. B. CT detector module with scintillator, electronics, amplification circuits, digitizer, and heatsink. image F10-19

      • (iv) Each ceramic detector element has a contact photodiode to generate charge from light.

      • (v) The detectors are designed in modules with electronics to acquire the digital data.

    • f. Emerging photon-counting detectors use solid-state detector technology.

      • (i) Energy discrimination allows individual photons to be assigned a discrete energy bin.

      • (ii) Multispectral imaging is possible without multiple scans at different tube potentials.

      • (iii) Technical challenge is the bandwidth of the electronics to measure high photon flux.

  • 6. Over beaming and geometrical efficiency

    • a. Overbeaming results from the presence of penumbra due to the finite focal spot size (Fig. 10-17).

    • b. Collimation extends beyond the active detector array configuration to ensure uniform beam.

      • (i) Active detectors in the penumbra region will generate a skewed slice sensitivity profile.

      • (ii) Manufacturers must increase the beam collimation beyond the outer active detectors.

      FIGURE 10-17 The x-ray beam is confined by the x-ray collimator and passes through the isocenter to the x-ray detector arrays. The detector width determines the spatial resolution in the z-dimension, while the beam width determines the coverage of the CT system along the z-axis of the patient. Note that the x-ray beam penumbra is outside of the active detector area. image F10-20

      FIGURE 10-18 The x-ray beam is loosely collimated so that the penumbra is positioned outside the active detector array. The shape of the slice sensitivity profile (SSP) for each detector array is uniform, but the SSP of the penumbra is anisotropic and is thus positioned just off the active detector arrays. This leads to higher dose levels to the patient. image F10-21

    • c. Beam geometrical efficiency is a function of the active detector width to beam width (Fig. 10-18).

      • (i) The penumbra width remains constant, independent of beam collimation.

      • (ii) For wider active detector arrays, efficiency is improved.

      • (iii) Smaller active detector widths (e.g., 8 channels of a 64 channel array) have poorer efficiency.

      • (iv) Dose due to overbeaming can be a large fraction of the total dose when a small number of detector arrays are active—low-dose efficiency triggers a warning on the CT scanner.

  • 7. Adapting data acquisition to patient anatomy: noise propagation in CT images

    • a. Noise in a pixel is the consequence of all projection data intersecting the voxel representing the pixel.

    • b. Noise variance (Σ2) within the pixel results from the propagation of noise variance of the projections:

    • c. “Noise adding in quadrature” means larger subcomponents of noise contribute more to the total noise.

      • (i) Total noise (the standard deviation) is the square root of σ2CT image.

    • d. Methods to reduce noise are focused on the components of the total noise that have greater impact.

      • (i) Bowtie filters modify the incident beam to achieve a more uniform x-ray transmission from the center to the periphery of the patient.

      • (ii) Tube current modulation varies the incident radiation as the x-ray tube rotates around the patient and adjusts current (x-ray output) on the attenuation path for each projection.

      • (iii) These issues are discussed next and later in the guide.

  • 8. Beam shaping filters

    • a. Most body parts are circular or approximately circular in shape.

    • b. Transmission of x-rays through the patient head or torso is greater in the periphery of the projection.

      FIGURE 10-19 A. The cross-section of most patients tends to be circular (bodies and heads), and in the absence of a beam shaping filter leads to the transmitted x-ray distribution with high levels at the periphery and low levels under the center of the patient. B. A beam shaping filter is positioned near the x-ray source to reduce the high x-ray fluence levels at the periphery of the fan beam relative to the center. The bow tie filter reduces the imbalance of signal levels received, and in doing so provides considerable radiation dose reduction to the periphery of the patient—with no loss in image quality. image F10-22

    • c. The bowtie filter shapes the incident radiation fluence to adjust for the attenuation variability presented by the patient (Fig. 10-19).

    • d. Reducing the beam intensity peripherally reduces dose to the patient with no appreciable loss of quality.

      FIGURE 10-20 The relative dose distributions in the circular patient cross-section with (A) no bow tie, (B) a well-designed bow tie filter, and (C) a bow tie filter that is too centrally focused. The primary role of the bow tie filter is to reduce the dose to the periphery of the patient. image F10-23

    • e. In the absence of a bowtie filter, the dose to the periphery is increased (Fig. 10-20A).

    • f. With the wrong bowtie filter mismatched to patient anatomy, dose increases centrally (Fig. 10-20C).

  • 9. View sampling

    • a. The data acquisition system (DAS) is intrinsic to the view sampling rate and sampling size.

    • b. The rotation period of the gantry about 360° occurs with continuous radiation output.

      • (i) View sampling rate (frequency) over the rotation period determines sampling size (Fig. 10-21).

      • (ii) A rate of 2,000 to 3,000 samples/s per detector element is necessary to maintain resolution.

    • c. View sampling size increases as r dθ (Fig. 10-21B), for angular sampling of .

      • (i) Spatial resolution decreases from the center to the edge of the FOV.

      FIGURE 10-21 A. Most CT scanners operate the x-ray tube in continuously on mode during CT acquisition, that is, the x-ray source is not pulsed. View sampling is defined by the data acquisition system associated with the detector arrays, which operate with relatively high bandwidth. For example, to acquire 1,000 views over 360° rotation of the gantry in 0.35 s, each detector needs to be sampled at almost 3,000 samples per second. B. This figure shows that for a given view sampling width/angle, the width of the sampling sector increases from the center of the field-ofview to the periphery, as r dθ. image F10-24

    • d. Focal spot dithering (Fig. 10-14) can help mitigate the loss of peripheral detail.

  • 10. High-resolution CT scanners

    • a. Specialty, small FOV scanners have been developed for years (small animal, specimen scanners).

      • (i) “Micro-CT” can deliver voxel dimensions on the order of 10 µm (but 30 minute scan time…).

    • b. Cone-beam CT scanners have high resolution because of small detector elements in the flat panel.

    • c. Whole-body CT scanners are now available to provide high spatial resolution in humans.

      • (i) One vendor provides 0.25 mm detector widths and 0.25 × 0.25 mm in-plane sampling.

      • (ii) Focal spots can limit spatial resolution, so six focal spots are provided from very small to large.

      • (iii) Reconstructed image matrix sizes for these scanners are 1,024 × 1,024 and 2,048 × 2,048.

    • d. Image noise is a challenge at high resolution—innovative reconstruction algorithms using artificial intelligence and deep learning algorithms may be a solution.

    • e. High-resolution CT may improve CT diagnostic accuracy across an array of clinical applications.


  • 1. The CT scan radiograph

    • a. X-ray tube and detector array are stationary—table moves through gantry with x-rays on.

      • (i) Scanned projection radiograph localizer also known as “scout” “topogram” “scanogram” …etc.

      • (ii) Technique factors use variable kV (often the same as the axial acquisitions) and low mA.

    • b. X-ray tube position at 0° (top), 90° (right), 180° (bottom), or 270° (left).

      • (i) Depending on patient position (supine/prone), anterior/posterior (AP) or posterior/anterior (PA) projections are acquired at 0° or 180°, and lateral projections at 90° or 270°.

      • (ii) Acquisition of both PA or AP and LAT localizer radiographs are common practice.

      • (iii) Lateral projection identifies positioning of the patient with the isocenter of rotation (table height).

      • (iv) PA acquisition is preferred to keep breast dose lower than AP acquisition for female patients.

    • c. The localizer defines anatomy slightly beyond the beginning and end of the planned CT acquisition.

      • (i) Using CT console software, the technologist plans the CT scan locations (Fig. 10-22).

    FIGURE 10-22 Most CT examinations start with the acquisition of the localizer, also known as the scan projection radiograph. This figure shows both the AP and lateral localizer images, which require separate scans. These images are acquired by keeping the gantry stationary (no rotation) while translating the patient on the table through the gantry. The scan projection radiograph is used by the technologists to define the position of the CT scan with respect to the patient’s anatomy, including stop and start points in z, and to adjust different regions of the scan for different acquisition parameters or post-acquisition parameters. image F10-25

  • 2. Axial (sequential) acquisition

    • a. Basic “step and shoot” where the tube and detector rotate around stationary patient/table (Fig. 10-23).

    • b. X-rays turn on for 360° to acquire raw data for a number of CT slices, then turn off.

    • c. Table is then moved with beam off to the next position and axial scan is repeated with stationary table.

    • d. Between each acquisition, the table moves a distance D, equal to the sampling distance at isocenter.

      • (i) For contiguous data acquisition D = nT, resulting in contiguous CT images along the z-direction

      • (ii) n = number of array channels and T = detector width; nT = collimated beam width + overbeaming

    • e. In some protocols, nT < D, which provides better sampling in z (e.g., for 3D volume rendering), but increases dose if the same technique factors are used.

  • 3. Helical (spiral) acquisition

    • a. Table moves simultaneous to tube and detector rotation, creating a helical path (Fig. 10-24).

    • b. X-rays are on constantly; eliminating table start/stop—removes intertial constraints.

    • c. The pitch describes the relative advance of the CT table per 360° rotation

      where Ftable is the table feed distance per 360°

      FIGURE 10-23 Axial (sequential) CT acquisition. With a stationary table, one data set is acquired over one rotation with x-rays on then off. The table translates a distance equal to the width of active detector arrays for contiguous scans, and the sequence is repeated to acquire the full dataset. image F10-26

      FIGURE 10-24 Helical (or spiral) CT is performed with continuous rotation of the gantry while the table is moving. Helical CT was made possible when slip rings were introduced to CT. With the gantry rotating, once the table is up to speed, helical CT scans proceed with no inertial impediment until the scan is completed. image F10-27

      Example: 40 mm = nT and 360° rotation time = 0.5 s; pitch = 1 occurs when Ftable = 80 mm/s

    • d. Pitch = 1 corresponds to contiguous axial scanning in principle.

    • e. Pitch >1 represents underscanning of some regions of the body.

      • (i) Allows for faster scanning (shorter total scan time) to reduce patient motion (e.g., pediatrics)

      • (ii) For technique factors held constant (except pitch), then

    • f. Pitch < 1 represents overscanning of the same regions of the body (and higher dose).

      • (i) Useful for multiplanar reformatting (sagittal, coronal) and 3D volume rendering.

      • (ii) Often, the technique factors (mAs) are reduced such that the dose is kept constant.

    • g. Comparison of axial, low pitch helical, and high pitch helical coverage (Fig. 10-25).

    • h. Acquired data at the beginning and end of scans do not have sufficient angular coverage to reconstruct artifact-free CT images (Fig. 10-26A), requiring over-Useful for multiplanar reformatting ranging.

      • (i) Results in wasted dose when the x-ray tube is on, equal to ½ nT at beginning and ½ nT at end

    • i. Adaptive beam collimation on some CT scanners eliminates the overranging penalty (Fig. 10-26B).

    FIGURE 10-25 The trajectory of the x-ray source around the patient is shown for (A) axial (sequential) imaging, (B) low pitch helical (spiral) imaging, and (C) high pitch helical imaging. image F10-28

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    Apr 18, 2023 | Posted by in GENERAL RADIOLOGY | Comments Off on Computed Tomography

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