DOI: 10.1148/rg.276075120
RadioGraphics 2007;27:1829-1837
© RSNA, 2007
AAPM/RSNA Physics Tutorials
AAPM/RSNA Physics Tutorial for Residents
Technologic Advances in Multidetector CT with a Focus on Cardiac Imaging1
Dianna D. Cody, PhD and
Mahadevappa Mahesh, PhD
1 From the Department of Imaging Physics, University of Texas M. D. Anderson Cancer Center, 1515 Holcombe Blvd, Unit 56, Houston, TX 77030 (D.D.C.); and the Russell H. Morgan Department of Radiology and Radiological Science, Johns Hopkins University School of Medicine, Baltimore, Md (M.M.). From the AAPM/RSNA Physics Tutorial at the 2005 RSNA Annual Meeting. Received May 17, 2007; revision requested June 20 and received July 6; accepted July 10. D.D.C. is a speaker for the Medical Technology Management Institute; M.M. receives research support from Siemens.
Address correspondence to D.D.C. (e-mail: dcody{at}mdanderson.org).
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Abstract
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Cardiac computed tomography (CT) is emerging as an important tool for the diagnosis and monitoring of heart disease. The prevalence of heart disease in the United States is already quite high and is expected to increase as the "baby boomer" segment of the population ages. To use complex multiple-row detector CT scanners most efficiently for cardiac examinations, it is important to understand many of the technical components. New developments in CT technology provide the ability to examine the structure of the heart with a level of detail that was not previously possible. In general, detector configurations have improved, the number of channels has increased, and rotation speed has increased, resulting in better quality of cardiac images. However, radiation dose for cardiac CT is fairly high and demands constant vigilance. Several steps can be taken to reduce the dose, including lowering the tube current as the x-ray beam crosses over certain areas of the body, decreasing the tube current during certain phases of the cardiac cycle, and using a higher pitch. Cardiac CT examination dose (for a coronary artery study) is approximately equivalent to that of an abdominal-pelvic CT examination or a dual-phase chest CT examination.
© RSNA, 2007
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Introduction
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Computed tomography (CT) scanners have been developing remarkably since around the year 2000. Substantial developments in hardware have resulted in dramatic improvements in scanning speed and a concurrent dramatic increase in the use of CT scanners for medical applications as a diagnostic tool. To use a state-of-the-art CT scanner efficiently, the operational team must understand how the technology functions. Such insight can lead to substantial improvements in patient throughput and image quality.
The recent major developments in CT scanner hardware are centered on the x-ray detectors, namely implementing multiple rows of detectors in the z direction. Although the materials used are not vastly different from those of previous scanners, the detector surface has undergone rapid change and has been carved up into small elements along the z-axis (patient table) direction. These detectors have always had separations along the other two directions, the x-y plane, in order to formulate the projection data used for image reconstruction. It is the separation of the detector surface along the orthogonal, or z-axis, direction that was completely new.
In addition to the new detector design, the electronic connections to the detector elements are also quite different from those of previous scanners. In previous conventional helical CT scanners, a single connection between each individual detector provided a single set of projection data for each rotation of the gantry assembly. In so-called multiple-row detector CT scanners, multiple connections between detector elements spread out along the z-axis direction provide multiple sets of projection data for each rotation of the scanner gantry assembly. The number of electronic connections available depends entirely on the vintage of the scanner model. As few as four of these electronic connections, also called data acquisition system channels, or as many as 256 can be found on currently installed multiple-row detector CT scanners.
In this article, we examine alternative detector array designs currently in use. In addition, we review the definition of pitch in the context of multiple-row detector CT and also examine radiation dose, with particular emphasis on dose delivered by cardiac CT examinations.
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Detector Configurations
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The detector array design in multiple-row detector CT varies with each manufacturer. The original four-channel scanners used several patterns of detector elements. The eight-channel scanners were constructed by using essentially the same detectors as those in the four-channel scanners. However, with the advent of 16-channel CT scanners, all CT manufacturers adopted a fairly similar design for their x-ray detectors. CT manufacturers also developed similar detector pattern designs for 32-, 40-, and 64-channel detector systems. In this section, we outline the similarities and differences among the CT manufacturers with respect to x-ray detector designs.
Uniform or Matrix Detectors
A detector design that is subdivided into equal elements, or portions, is called "uniform," "matrix," or "mosaic." GE Healthcare (Waukesha, Wis), the first CT manufacturer to bring a multiple-row detector CT scanner design to the market, used this design in their four- and eight-channel scanners (Fig 1). It has also been adapted for use in several more recent scanner designs (ie, 40-and 64-channel scanners).
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Figure 1. Mosaic detectors have elements that are all a uniform size. The thickness of the sections that can be generated from these detectors is a multiple value of the uniform size of the detector element. (In this case, sections can have a thickness of 1.25, 2.5, 3.75, 5, 7.5, or 10 mm.)
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Nonuniform or Variable Detectors
Several CT manufacturers used a different design to carve the detector surface into smaller elements in the orthogonal direction or along the z axis (the longitudinal axis of the patient table). In this design, the element sizes are variable and not uniform (Fig 2). In combination with a physical postpatient collimator, signals from these detectors can be combined to form various values of image thickness in a manner that improves detector efficiency over that of uniform detectors. This improvement is a result of the smaller proportion of dead space due to the divisions (cell walls) between detector elements. Although several CT manufacturers (eg, Siemens [Erlangen, Germany] and Philips Medical Systems [Eindhoven, the Netherlands]) used this approach for their early four-channel CT scanner designs, it has not since been incorporated into later-generation multiple-row detector CT scanners.
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Figure 2. Variable or nonuniform detectors are composed of elements that are not uniform in size but can be combined with a postpatient collimator unit to generate sections with several different thickness values. (In this case, sections can have a thickness of 1, 2.5, 5, or 10 mm.)
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Hybrid Detectors
Toshiba (Nasu, Japan) was the first to introduce the combination of two discrete detector element sizes in its four-channel CT scanner (and subsequently in its eight-, 16-, 32-, and 64-channel scanners). This detector design has a number of narrow detector elements in the center of the detector and a different number of wider detectors (usually double the width of the narrow detectors) on both sides of the span of narrow detectors (Fig 3). The number of narrow and wider detectors can vary. Current 16-channel scanner models all use a hybrid detector design pattern.
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Figure 3. Hybrid detectors are composed of elements that are not uniform in size, but in general only a limited number of element sizes is used. This design has been especially popular on 16-channel CT scanners across all manufacturers. The detector design shown is from a Toshiba four-channel scanner.
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Four-Channel to 64-Channel Options
In conventional single-channel helical CT scanners, a single connection between each individual detector provided a single set of projection data for each rotation of the gantry assembly. 
In multiple-row detector CT scanners, multiple connections (or channels) between detector elements that are spread out along the z axis provide multiple sets of projection data for each rotation of the scanner gantry assembly. The number of electronic connections available depends entirely on the model of the scanner, although there has been a general tendency toward more channels over time.
These channels are no longer fixed but are instead variable, so that they can sample different detector elements and can even sample several detector elements simultaneously, effectively adding (or binning) their signals together. The concept of variable detector sampling is new to CT scanner designs, and it provides flexibility for image reconstruction options that were unavailable in conventional helical CT scanners.
CT scanners can be purchased with various numbers of channels: four-, six-, 10-, 16-, 32-, 40-, and 64-channel systems are now commercially available (Fig 4). Some manufacturers equip their CT scanners with a standard minimum rotation speed and also offer optional upgrades for faster rotation speeds (eg, to support cardiac CT capability). In February 2007, a prototype CT scanner with 256 channels was installed for clinical use at Johns Hopkins Hospital with a coverage of 12.8 cm at the isocenter per gantry rotation. The goal is to scan the entire cardiac region in a single gantry rotation. The "slice war" appears to be continuing with little slowdown in sight.
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Definition of Pitch
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Pitch is defined as the ratio of the distance the patient table travels during one rotation of the x-ray beam to the total x-ray beam width. To understand pitch, think of the CT x-ray beam as spray paint coming out of the can. If the table were to advance exactly in concert with the gantry spin, that is, the width of the paint spray, the patient would be painted uniformly. If the table were to move a bit faster than the gantry spin, the patient would have a candy cane–striped appearance with unmarked flesh between ribbons of paint. If the table were to move a bit slower than the gantry spin, the patient would be coated with ribbons of overlapping paint, and the color would be darker on the overlapping areas than elsewhere.
In the first case (uniform color), the pitch would be equal to unity. In the second case (candy cane color), the pitch would be greater than 1, and in the third case (light and dark ribbons of color), it would be less than 1. The above definition of pitch remains consistent with single-row detector CT scanners. In multiple-row detector CT scans, however, the x-ray beam width can be more difficult to determine. It is calculated by multiplying the number of channels active during the image acquisition (this number may not be equal to the total number of available channels) by the width of the detector surface assigned to each data acquisition system channel during the image acquisition (channel width).
The flexibility of multiple-row detector CT resides in its ability to assign a different number of detector elements to each data channel, so that the sampling of the attenuated path through the patient can be customized for different imaging tasks. This flexibility can make the computation of pitch somewhat confusing. If the detector configuration (the number of active channels and the detector width of each channel) is clearly shown on the CT scanner console, the nominal overall beam width can be found by multiplying the number of active channels by the channel width. For example, a detector with 64 channels and a 0.625-mm channel width requires a 40-mm beam width (Fig 4).
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Review of CT Dose
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Definitions
The ability of x-rays to ionize air is known as exposure. Exposure generally refers to a measurement at some specific point in space and is expressed in units of roentgens (coulombs per kilogram). This quantity indicates how much radiation is present but not how much is absorbed.
The amount of energy that is actually absorbed at a specific point in space is known as the absorbed dose, and it is expressed in units of rads (1 rad = 100 erg/g) or grays (1 Gy = 1 J/kg). This quantity indicates how much radiation is absorbed but not where it is absorbed or the risk to those tissues being irradiated. Absorbed dose can be estimated by using devices such as thermoluminescence dosimeters. These dosimeters are made of special materials; after being exposed to radiation, they can be heated so that they emit light in proportion to the energy (dose) deposited by the ionizing radiation. This type of dosimeter can be used as a point detector to assess absorbed dose in CT, but it has many disadvantages that make it less useful for routine dosimetry.
Absorbed dose is used to estimate the effective dose, which takes into account where the radiation is being absorbed and the radiosensitivity of the tissues absorbing the radiation. The effective dose depends on two factors; one is the tissue type being irradiated (relatively sensitive or insensitive to radiation damage), and the other is the type of radiation being delivered (the radiation weighting coefficient, currently equal to unity for x-rays). The unit for effective dose is the rem or the sievert, and the effective dose provides a means for converting the dose administered during an x-ray examination of a limited part of the body (such as a cardiac CT study) to the equivalent dose as if the entire body had been exposed. Calculation of the effective dose also allows us to compare radiation dose values delivered by different types of radiographic examinations and to the average annual background dose delivered to the U.S. population (3.6 mSv/y).
Dosimetry Phantoms
In CT, a CT dose index (CTDI) value is calculated to represent the radiation dose delivered during a CT examination (1). Standard cylindrical acrylic dose phantoms (16 or 32 cm in diameter and 15 cm long) are used to measure CT dose. These phantoms are fitted with holes to allow CT ion chambers (approximately 1 cm in diameter with a 10-cm active length) to be easily inserted. A central hole is available in addition to holes at four orthogonal peripheral positions (generally referred to as the 12-, 3-, 6-, and 9-o’clock positions), each 1 cm below the surface. The smaller phantom is used to assess the dose for head CT examinations and pediatric body CT examinations; the larger phantom is used to assess the dose for adult body CT examinations.
During a CT examination, the energy deposited in an attenuating material (such as air or an acrylic phantom) during a single rotation of the x-ray tube can be depicted along the z axis (the length of the patient table), as shown in Figure 5. This depiction is often called a dose profile and can be measured with a CT ion chamber, typically 10 cm (or 100 mm) long. The energy deposited in the dose profile is called CTDI100 (to indicate the length of the ion chamber in millimeters) and is measured in milligrays. The area under the profile curve (Fig 5) is mathematically equal to the dose delivered during a single axial rotation of the x-ray tube around the ion chamber inside the acrylic CTDI phantom.
Key Dose-related Parameters in CT
The dose profile is a useful concept, but alone it is not sufficient to adequately describe the radiation dose delivered by a CT examination.
Weighted CTDI.—
When an x-ray beam is applied to a patient, the beam often passes directly into the entire surface of the patient’s body. The direct, or primary, x-ray beam contains the most photons and is most attenuated by the patient tissue. Therefore, the energy delivered to the patient is higher at the skin surface than in the center of the body, even though the x-ray beams from all the way around the gantry pass through the center of the body. The fact that the surface is exposed to more photons than the center portion requires an adjustment to CTDI100 to account for this nonuniform dose delivery. This new CTDI value is called the weighted CTDI (CTDIw), and to calculate it, we multiply the peripheral CTDI100 value by two-thirds and the center CTDI100 value by one-third.
Volume CTDI.—
Nearly all current CT scanners can be operated in helical mode, in which the patient table advances smoothly as the gantry spins. We adjust for pitch (the gap or overlap of the helical pattern of radiation) by calculating what is known as volume CTDI (CTDIvol), which is CTDIw divided by the pitch. A scan with a pitch less than unity would therefore have a CTDIvol value larger than the CTDIw value, and scans with a pitch greater than unity would have a CTDIvol value smaller than the CTDIw value.
Dose-Length Product.—
The CTDIvol value and dose-length product (DLP) are sometimes displayed on the scanner console during the CT examination. The DLP is calculated to account for the differences in the scan extent for a CT examination. For some patients, we perform only a limited examination (eg, a small portion of the chest to follow up an individual nodule); for others, we scan a large extent (eg, from the chest to the toes for a vascular runoff examination). The risk of radiation-induced damage to patient tissues increases with the volume of body scanned. The DLP is calculated by multiplying CTDIvol in grays by the scan extent in centimeters. The DLP essentially provides a value that allows us to compute the desired quantity, the effective dose.
Effective Dose.—
Effective dose is not calculated automatically by current commercial scanners. We can compare the effective dose with that of other radiographic examinations and with annual background dose. Understanding this term should help us understand why CT dose estimates cannot be readily customized for individual patients.
To assess the risk of radiation for an individual patient, we need to know exactly which organs have been irradiated. Sometimes this information is available, but it can often be difficult to determine how much of the thyroid, for example, was in the x-ray beam for a typical chest examination. In some patients, sensitive organs are nearer to the surface than they are in other patients. We know that tissue closer to the skin attenuates more photons and so would be at a higher risk of damage. If oral contrast material is present in the gut, it can cause a different pattern of attenuation than if it is not. The number and variable effects of all these possible complicating factors prohibit today’s technology from making precise estimates of the dose delivered to an individual patient during a CT examination.
Thus, to estimate the effective dose to a patient, we use a rough approximation, multiplying the CTDIvol by a constant value (called the k factor) that generally accounts for the sensitive organs that may be affected by the x-ray beam during CT (2). This constant is expressed in units of millisieverts per milligray-centimeters, yielding an effective dose estimate in millisieverts. The k factor used to estimate the effective dose in cardiac CT is that used for the chest, 0.017 mSv/ mGy-cm (2). Other k factors should be used for CT examinations that cover the head, neck, abdomen, and pelvis regions.
In summary, we use index-based values to describe dose in CT, modified by a series of adjustments to reflect the dose of radiation that the scanner delivers to a plastic phantom. Although customized dose calculations for individual patients theoretically can be generated by using data present in the CT images themselves, no automated method has been developed that can recognize and segment the various tissues and organs for individual patients. Another limitation of the CTDI approach is its inapplicability to true cone-beam CT systems that use large-area radiation detectors, such as those attached to radiation treatment devices, and to rotating C-arm systems, such as those gaining popularity in interventional radiology settings.
Dose Efficiency in Multidetector CT
Dose efficiency (also called geometric efficiency) is of particular concern with the earlier four-channel multiple-row detector CT scanners, for which the x-ray photon beam has to be quite uniform as it strikes the detector array. This requirement means that the natural shadowing of the beam (penumbra) attributable to the finite-sized focal spot is intentionally positioned to strike the neighboring nonactive detector elements. Thus, radiation is passed through the patient that does not contribute to image generation (called overbeaming). The width of the penumbra is fairly constant with each scanner, generally in the range of 1–3 mm.
The proportion of radiation wasted relative to the overall width of the x-ray beam varies with the protocol used. If thin sections are required and the overall x-ray beam width is small—5 mm, for example—then the proportion of wasted x-rays could be 20%–60% (resulting in a dose efficiency of 40%–80%). If thin sections are not required and a wider x-ray beam can be used—20 mm, for example—then the proportion of wasted x-rays would be 5%–15% (resulting in a dose efficiency of 85%–95%). Dose efficiency may be displayed on the scanner console. More recent multisection CT scanners have been engineered such that this problem is very much diminished.
As shown in Figure 6a, the penumbra effect can represent a large percentage of the overall x-ray beam, which needs to cover all the active detectors. This is the case for protocols requiring few thin detectors, for example 4 x 0.625 mm or 4 x 0.5 mm. Conversely, in Figure 6b, even though there is overbeaming, the penumbra effect is smaller compared to the primary x-ray beam. Therefore, dose efficiency is less of a concern for CT scanners with 16 channels and greater (the beam width can remain much larger even when collecting finely spaced projection data).
The effect of dose efficiency is demonstrated in Figure 7, where normalized values of CTDI are shown among four-, 16-, and 64-channel multiple-row detector CT scanners from the same manufacturer. The CTDIw (normalized to 16- section CT values) shows significantly higher results for four-channel scanners compared with 16- and 64-channel scanners. Dose increases markedly with narrowing beams and fewer active detector elements, but less so for wider beams and more detectors owing to improved geometric efficiency.
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Figure 7. CTDIw values (normalized to the 16-channel scanner values) for four- (left), 16- (center), and 64- (right) channel multiple-row detector CT scanners from the same manufacturer. The four-channel acquisitions deliver higher dose index values than the 16- and 64-channel acquisitions owing to poor dose efficiency. Red bars = head CT dose, yellow bars = body CT dose.
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New Scanner Design: Dual-Source Cardiac CT Scanner
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In 2006, Siemens released a new CT scanner with two x-ray tubes and two sets of x-ray detectors. This new scanner, called the Somatom Definition, is particularly well suited to performing cardiac CT scans. Because the minimum rotation speed required for a cardiac CT image is equal to just over one-half of a rotation for a conventional single-source multisection CT scanner, it is also equal to just over one-fourth of a rotation for a dual-source multisection CT scanner. Thus, this new design essentially improves the temporal resolution provided for cardiac CT by a factor of two (3). In addition, linking the patient’s heart rate to the pitch (by which a higher rate can be examined by using a higher pitch) effectively reduces the dose of radiation delivered during the examination and can also eliminate the need for additional drugs (eg, beta-blockers) before the cardiac CT examination to reduce the heart rate.
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Cardiac CT Dose
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Cardiac CT is emerging as an important tool for the diagnosis and monitoring of heart disease. The incidence of heart disease in the United States is already quite high and is expected to increase as the "baby boomer" segment of the population ages. To use complex multiple-row detector CT scanners most efficiently for cardiac examinations, it is important to understand many of the technical components. In addition, radiation dose for cardiac CT studies is fairly high and demands constant vigilance.
Cardiac CT techniques are generally similar to those used for routine chest CT—with a few important exceptions. Although the tube potential (measured in kilovolts) and current (measured in milliamperes) are often similar to those used for a routine chest CT study, the rotation speed is generally the fastest that can be achieved by the scanner model, and the pitch value used is often much lower than that used in routine chest examinations (typically 0.2–0.4). The use of extremely low pitch values can cause some alarm with respect to radiation dose values. However, the scan extent (start-to-stop distance for the image acquisition) is generally limited (usually 12 cm), which results in a dose range that is near accepted values for similar CT examinations of the torso.
Typical Cardiac CT Dose Values
Some recently reported effective dose values for cardiac CT examinations are shown in Table 1. Historically, CT facilities have chosen to implement protocols specific to the preference of their interpreting radiologists, which has meant that a wide variety of parameter settings have been used to achieve the same desired result. The case is no different for cardiac CT—different facilities have implemented different protocols for their cardiac CT examinations. The range of dose values shown in Table 1 is a result of the variations in techniques and CT scanners between facilities. Obviously, a calcium scoring examination requires a much smaller radiation dose than a CT arteriogram. What should be noted is that the dose values obtained by using thermoluminescence detectors and the dose values obtained by using the CTDI method overlap considerably, increasing our confidence in the assessed dose, no matter which detector type is used.
Typical cardiac CT dose values are compared with those of other routine CT examinations in Table 2 (4–8). Although cardiac CT dose does appear to be on the higher end of the CT radiation dose spectrum, it is quite similar to the dose given in a routine combined abdomen and pelvis CT examination or a dual-phase chest CT examination (with and without intravenous contrast material).
Reduction of Cardiac CT Dose
The dose of radiation delivered during a cardiac CT examination can be decreased by implementing features such as tube current modulation, cardiac phase–specific scanning, and increased pitch. For instance, all scanners are designed so that the x-ray tube current can be modulated, or varied, while the gantry spins around the patient. The tube current thus can be decreased when the beam crosses areas of the body with less tissue or lower tissue density, thereby lowering the total overall dose of radiation needed (9). In general CT examinations, dose reductions on the order of 20%–40% have been achieved with these approaches (10–12).
A dose-reduction approach specific to cardiac CT is to decrease the x-ray tube current during phases of the cardiac cycle that are expected to contain fairly large amounts of motion. Images acquired during these phases are expected to be of lower quality and of less value for interpretation; thus, the radiologist may choose that they not be captured for use in image reconstruction. Dose reductions of up to 50% have been achieved with this feature (13). This option may or may not be successful at individual facilities; many radiologists appear to demand images from all cardiac cycles to review regardless of their temporal resolution quality.
A third means of reducing the effective dose is to use a higher pitch value. This method is particularly used for selective patients with higher-than-normal heart rates, as with the new dual x-ray tube CT scanner discussed earlier (14). The overall dose of radiation delivered can also be substantially decreased with this method.
Combining two or more of these dose reduction options during a cardiac CT study may result in even greater decreases in the dose of radiation delivered.
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Summary
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The new developments in technology for CT provide the ability to examine the heart structure with a level of detail that was not previously possible. In general, detector configurations have been improved, the number of channels has generally increased, and rotation speed has increased, resulting in better quality of cardiac images. Radiation dose becomes an issue for cardiac CT primarily due to low pitch values required for the special reconstruction process. However, several steps can be taken to reduce the dose, including lowering the tube current as the x-ray beam crosses over certain areas of the body, decreasing the x-ray tube current during certain phases of the cardiac cycle, and using a higher pitch value. Cardiac CT examination dose (for a coronary artery study) is approximately equivalent to that of an abdomen-pelvis CT examination or a dual-phase chest CT examination.
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Footnotes
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Abbreviations: CTDI = CT dose index
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References
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- European guidelines on quality criteria for computed tomography [EUR 16262 EN] 1999. http://www.drs.dk/guidelines/ct/quality/index.htm. Accessed July 3, 2007.
- McCollough CH, Schmidt B, Primak A, et al. Measurement of temporal resolution in dual source CT [abstr]. In: Radiological Society of North America scientific assembly and annual meeting program. Oak Brook, Ill: Radiological Society of North America, 2006; 463.
- Hunold P, Vogt FM, Schmermund A, et al. Radiation exposure during cardiac CT: effective doses at multi–detector row CT and electron-beam CT. Radiology 2003;226:145–152.[Abstract/Free Full Text]
- Trabold T, Buchgeister M, Kuttner A, et al. Estimation of radiation exposure in 16-detector row computed tomography of the heart with retrospective ECG-gating. Rofo 2003;175:1051–1055.[Medline]
- McCollough CH. Patient dose in cardiac computed tomography. Herz 2003;28:1–6.[CrossRef][Medline]
- Bae KT, Hong C, Whiting BR. Radiation dose in multidetector row computed tomography cardiac imaging. J Magn Reson Imaging 2004;19:859–863.[CrossRef][Medline]
- Gerber TC, Kuzo RS, Morin RL. Techniques and parameters for estimating radiation exposure and dose in cardiac computed tomography. Int J Cardiovasc Imaging 2005;21:165–176.[CrossRef][Medline]
- McCollough CH, Bruesewitz MR, Kofler JM. CT dose reduction and dose management tools: overview of available options. RadioGraphics 2006;26: 503–512.[Abstract/Free Full Text]
- Kalra MK, Maher MM, Kamath RS, et al. Sixteen–detector row CT of abdomen and pelvis: study for optimization of z-axis modulation technique performed in 153 patients. Radiology 2004; 233:241–249.[Abstract/Free Full Text]
- Graser A, Wintersperger BJ, Suess C, Reiser MF, Becker CR. Dose reduction and image quality in MDCT colonography using tube current modulation. AJR Am J Roentgenol 2006;187:695–701.[Abstract/Free Full Text]
- Rizzo S, Kalra M, Schmidt B, et al. Comparison of angular and combined automatic tube current modulation techniques with constant tube current CT of the abdomen and pelvis. AJR Am J Roentgenol 2006;186:673–679.[Abstract/Free Full Text]
- Jakobs TF, Becker CR, Ohnesorge B, et al. Multi-slice helical CT of the heart with retrospective ECG gating: reduction of radiation exposure by ECG-controlled tube current modulation. Eur Radiol 2002;12:1081–1086.[CrossRef][Medline]
- McCollough CH, Primak AN, Saba O, et al. Dose performance of a 64-channel dual-source CT scanner. Radiology 2007;243:775–784.[Abstract/Free Full Text]
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Abstract
Introduction
