Siemens Ct Scan 32 Slice
AbstractThis article describes the principles and evolution of multislice CT (MSCT), including conceptual differences associated withslice definition, cone beam effects, helical pitch, and helical scan technique. MSCT radiation dosimetry is described, anddose issues associated with MSCT—and with CT in general—as well as techniques for reducing patient radiation dose are discussed.Factors associated with the large volume of data associated with MSCT examinations are presented.Keywords.This article, the third in a series of continuing education articles on the principles of CT, focuses on multislice CT (MSCT). Limitations of Single-Slice Slip Ring and Helical ScannersSoon after their introduction in the late 1980s, slip ring scanners and helical (spiral) CT were rapidly adopted and soonbecame the de facto standard of care for body CT. However, a significant problem became evident: helical CT was very hardon x-ray tubes. For example, an abdomen–pelvis helical CT covering 60 cm (600 mm) of anatomy with a 5-mm slice thickness,a pitch of 1.0 (thus requiring 120 rotations), and typical technique factors (120 kilovolts peak kVp, 250 mA, 1-s rotationtime) deposits a total of 3.6 × 10 6 J of heat in the x-ray tube anode. Before slip ring CT, individual slices obtained with an equivalent technique (120 kVp,250 mA, 1-s scan) would deposit only 30,000 J, much of which could be dissipated during the relatively lengthy (several seconds)interscan delay.A limitation imposed by tube heating was that the thin slices (. MSCT DetectorsThe primary difference between single-slice CT (SSCT) and MSCT hardware is in the design of the detector arrays, as illustratedin.
SSCT detector arrays are one dimensional ; that is, they consist of a large number (typically 750 or more) of detector elements in a single row across the irradiatedslice to intercept the x-ray fanbeam. In the slice thickness direction ( z-direction), the detectors are monolithic, that is, single elements long enough (typically about 20 mm) to intercept the entirex-ray beam width, including part of the penumbra (here, the term “x-ray beam width” always refers to the size of the x-raybeam along the z-axis—that is, in the slice thickness direction). In MSCT, each of the individual, monolithic SSCT detector elements in thez-direction is divided into several smaller detector elements, forming a 2-dimensional array.
Siemens 32 Slice Ct Scan
Rather than a single row of detectors encompassing the fanbeam, there are now multiple, parallel rows of detectors. Before further discussion, a comment on nomenclature is called for: the use of the term “MSCT” is not universal. Others usethe terms “multirow CT” and “multidetector row CT (MDCT)” because they are more descriptive of this technology than the term“multislice CT.” Throughout this article, however, the term “MSCT” is used.The first scanner with more than one row of detectors and a widened z-axis x-ray beam was introduced by Elscint in 1992 (CT-Twin). This scanner had 2 rows of detectors, allowing data for 2 slicesto be acquired simultaneously, and was developed primarily to help address the x-ray tube heating problem. As a curious historicalnote, according to the description given earlier in this article, the first MSCT scanner would actually be the first-generationEMI Mark 1. With 2 adjacent detectors and a widened x-ray beam, this scanner collected data for 2 slices at the same timeand thereby reduced the lengthy examination time associated with the 5- to 6-min scan time ( ). The first scanners of the “modern MSCT era” were introduced in late 1998 and are described in the following discussion( ).
The costs of a used CT system in this range is typically 35-60.000 euros. Popular systems: Toshiba Asteion S4, Toshiba Asteion Multi, Toshiba Aquilion Multi, and GE Lightspeed Ultra. 16-32 Slice CT Scanners. The next category is the CT scanners with 16-32 slices. What is the price range? Typically 60-100.000 euros. CT Scan Protocols, CT Protocols by Manufacturer- GE, Siemens, Phillips, Toshiba. Slice Counts- Dual Source, 320 slice, 256 slice, 128 slice, 64 slice, 16 slice, 4 slice, single detector. All Sections Journal Club Lectures Pearls Protocols Case Studies Medical Illustrations Search. This feature is not available right now. Please try again later.
MSCT Data AcquisitionA detector design used in one of the first modern MSCT scanners consisted of 16 rows of detector elements, each 1.25 mm long in the z-direction, for a total z-axis length of 20 mm. Each of the 16 detector rows could, in principle, simultaneously collect data for 16 slices, each 1.25mm thick; however, this approach would require handling an enormous amount of data very quickly, because a typical scannermay acquire 1,000 views per rotation. If there are 800 detectors per row and 16 rows, then almost 13 million measurementsmust be made during a single rotation with a duration of as short as 0.5 s.Because of the initial limitations in acquiring and handling such large amounts of data, the first versions of modern MSCTscanners limited simultaneous data acquisition to 4 slices. Four detector “rows” corresponding to the 4 simultaneously collectedslices fed data into 4 parallel data “channels,” so that these 4-slice scanners were said to possess 4 data channels. These4-slice scanners, however, were quite flexible with regard to how detector rows could be configured; groups of detector elementsin the z-direction could be electronically linked to function as a single, longer detector, thus providing much flexibility in theslice thickness of the 4 acquired slices.
Examples of detector configurations used with the 4 channels are illustrated infor 2 versions of 4-slice MSCT detectors: one based on the detector design described earlier (16 rows of 1.25-mm elements)and the other based on an “adaptive array” consisting of detector elements of different sizes (other detector designs wereused by other manufacturers) (,). FIGURE 2.Flexible use of detectors in 4-slice MSCT scanners. (A) Groups of four 1.25-mm-wide elements are linked to act as 5-mm-widedetectors.
(B) Inner 8 elements are linked in pairs to act as 2.5-mm detectors. (C) Inner, adaptive-array elements are linkedto act as 5-mm detectors (1 + 1.5 + 2.5) and, together with outer, 5-mm elements, yield four 5-mm slices. (D) The 4 innermostelements are linked in pairs to form 2.5-mm detectors (1 + 1.5), which along with the two 2.5-mm detectors, collect data forfour 2.5-mm slices.
Possible detector configurations for the detector design encompassing 16 rows of 1.25-mm elements for the acquisition of 4slices are illustrated in. In, 4 elements in a group are linked to act as a single 5-mm detector (4 × 1.25). The result is four 5-mm detectors coveringa total z-axis length of 20 mm. When a 20-mm-wide x-ray beam is used, 4 slices with a thickness of 5 mm are acquired. The acquired5-mm slices can also be combined into 10-mm slices, if desired.
In, pairs of detector elements are linked to function as four 2.5-mm detectors (2 × 1.25). When a 10-mm-wide x-ray beam is used,four 2.5-mm slices can be acquired simultaneously. Again, the resulting 2.5-mm slices can be combined to form 5-mm slices(5-mm axial slices are generally preferred for interpretation purposes).
A third possibility is to use a 5-mm-wide x-ray beamto irradiate only the 4 innermost individual detector elements for the acquisition of four 1.25-mm slices. Yet another possibilityis to link elements in triplets and use a 15-mm-wide x-ray beam to acquire four 3.75-mm slices.Similarly, the individual elements of the adaptive array can be appropriately linked to acquire four 5-mm slices or four 2.5-mm slices. Another possibility is to use a 4-mm-wide x-ray beam (which would irradiate only part of the 1.5-mm elements) to yieldfour 1-mm slices. Thinner slices can be combined to form thicker slices for interpretation purposes, if necessary.As data acquisition technology advanced, more data channels were provided to allow the simultaneous acquisition of more than4 slices.
An 8-channel version of the system encompassing the detector array in (introduced approximately 3 y later) could acquire eight 2.5-mm slices or eight 1.25-mm slices (which could be combined toform thicker slices for interpretation). Submillimeter Slices and Isotropic ResolutionThe 4-slice and 8-slice MSCT scanners just described were also capable of acquiring ultrathin (“submillimeter”) slices (butonly 2 at a time) by collimating the x-ray beam in the z-axis to partially irradiate the 2 innermost detector elements in the detector array. For example, for the detector arrayin, if the x-ray beam is collimated to a 1.25-mm width and aligned so as to straddle and partially irradiate the 2 innermostdetector elements, then 2 slices, each 0.625 mm thick, can be obtained. When images resulting from such an acquisition arereformatted into sagittal, coronal, or other off-axis images, the reformatted images exhibit spatial resolution in the z-direction that is essentially equal to that within the plane of the axial slices.
Resolution that is (essentially) equalin all 3 directions is said to be isotropic.Because only 2 submillimeter slices could be acquired simultaneously with these earlier MSCT scanners, this capability wasnot widely used because of limited z-axis coverage and tube heating limitations. Submillimeter scanning had to await the introduction of 16-slice scanners. 16-Channel (16-Slice) Scanners—and MoreThe installation of MSCT scanners providing 16 data channels for 16 simultaneously acquired slices began in 2002. In additionto simultaneously acquiring up to 16 slices, the detector arrays associated with 16-slice scanners were redesigned to allowthinner slices to be obtained as well.
Detector arrays for various 16-slice scanner models are illustrated in. Note that in all of the models, the innermost 16 detector elements along the z-axis are half the size of the outermost elements, allowing the simultaneous acquisition of 16 thin slices (from 0.5 mm thickto 0.75 mm thick, depending on the model). When the inner detectors were used to acquire submillimeter slices, the total acquiredz-axis length and therefore the total width of the x-ray beam ranged from 8 mm for the Toshiba version to 12 mm for the Philipsand Siemens versions. Alternatively, the inner 16 elements could be linked in pairs for the acquisition of 16 thicker slices( ). During 2003 and 2004, MSCT manufacturers introduced models with both fewer than and more than 16 channels. Six-slice and 8-slicemodels were introduced by manufacturers as cost-effective alternatives.
At the same time, 32-slice and 40-slice scanners werebeing introduced.By 2005, 64-slice scanners were announced, and installations by most manufacturers began. Detector array designs used by severalmanufacturers are illustrated in. The approach used by most manufacturers for 64-slice detector array designs was to lengthen the arrays in the z-direction and provide all submillimeter detector elements: 64 × 0.625 mm (total z-axis length of 40 mm) for the Philips and GE Healthcare models and 64 × 0.5 mm (total z-axis length of 32 mm) for the Toshiba model. The design approach of Siemens was quite different. The detector array of theSiemens 32-slice scanner (containing 32 elements each 0.6 mm long, for a total z-axis length of 19.2 mm) was combined with a “dynamic-focus” x-ray tube for the simultaneous acquisition of 64 slices.
Thisx-ray tube could electronically—and very quickly—shift the focal spot location on the x-ray tube target so as to emit radiationfrom a slightly different position along the z-axis. Each of the 32 detector elements then collected 2 measurements (samples), separated along the z-axis by approximately 0.3 mm. The net result was a total of 64 measurements (32 detectors × 2 measurements per detector)along a 19.2-mm total z-axis field of view (this process is referred to in Siemens literature as “Z-Sharp” technology) ( ). In the preceding examples, in addition to the simultaneous acquisition of more slices, MSCT x-ray beam widths can be considerablywider than those for SSCT. Sixteen-slice MSCT beam widths are up to 32 mm; 64-slice beams can be up to 40 mm wide; and evenwider beams are used in systems currently under development or in clinical evaluation. A possible consequence is that morescatter may reach the detectors, compromising low-contrast detection.
Generally, however, the antiscatter septa traditionallyused with third-generation CT scanners can be made sufficiently deep to remain effective with MSCT. An example of a sectionof a 16-slice detector with the associated scatter removal septa is shown in. MSCT Slice Thickness and X-Ray Beam CollimationIn SSCT, slice thickness is determined by prepatient and possibly postpatient x-ray beam collimators. Generally, the x-raybeam collimation was designed such that the z-axis width of the x-ray beam at the isocenter (i.e., at the center of rotation) is the same as the desired slice thickness.(The x-ray beam width, usually defined as the full width at half maximum FWHM of the z-axis x-ray beam intensity profile, is discussed in detail in the second article in this series ( ).)In MSCT, however, slice thickness is determined by detector configuration and not x-ray beam collimation. For example, the4 slices in are each 5 mm thick because they are acquired by 5-mm detectors (formed by linking four 1.25-mm detector elements). The 4slices in are 2.5 mm thick because they are acquired by 2.5-mm detectors (formed by linking two 1.25-mm elements).
Siemens Ct Scan
Because it is thelength of the individual detector (or linked detector elements) acquiring data for each of the simultaneously acquired slicesthat limits the width of the x-ray beam contributing to that slice, this length is often referred to as detector collimation.In, the detector collimation is 5 mm. In, the detector collimation is 2.5 mm. The actual x-ray beam collimation is not directly involved in determining slice thickness,other than that the “total” z-axis beam collimation should be equal to the total thickness of the 4 slices, for example, 20 mm in or 10 mm in (that this is not necessarily true is discussed in the MSCT dosimetry section later in this article) ( ). Cone Beam Effects in MSCTCone beam effects in CT are associated with the divergent nature of the x-ray beam emitted from the x-ray tube. This divergencemeans that the z-axis of the x-ray beam is somewhat wider when it exits the patient than when it entered the patient.In SSCT, the main consequence of x-ray beam divergence is the potential for partial-volume streaking, discussed in the secondarticle in this series ( ) and reviewed here. During a 360° rotation, the same path (or nearly the same path) of x-rays through the patient is measuredtwice, but with x-rays traveling in opposite directions, for example, once with the tube above the patient and detectors belowand later on during the rotation with the tube below the patient and the detectors above (these are referred to as parallelopposed rays or conjugate rays). Because of beam divergence, however, the cone-shaped x-ray beam samples slightly differenttissue volumes in each direction , potentially leading to data inconsistencies and streak artifacts.
The thicker the slice (the wider the x-ray beam), themore pronounced the divergence and the more likely it is that parallel opposed ray measurements will be inconsistent. Cone beam effects are more severe in MSCT. Consider the MSCT configuration in, collecting data for eight 2.5-mm slices (total beam width of 20 mm).
Note that the measurements obtained with the outermostdetector (shading in ) from opposite sides of the patient not only sample different tissues but also do not even lie within the same slice ( ).With 4-slice scanners, the total x-ray beam width was sufficiently narrow (e.g., 5 mm wide for four 1.25-mm slices) or elsethe slices were sufficiently thick (four 5-mm slices) that cone beam effects were tolerable and conventional filtered backprojectionreconstruction was still usable. However, MSCT scanners of later generations, which collected more and thinner slices, requiredthe development of alternate cone beam reconstruction algorithms (which are beyond the scope of this article). Because ofcone beam effects, some MSCT scanners with 16 channels or more only allow the simultaneous acquisition of the maximum numberof slices (e.g., 16 slices by a 16-channel scanner) during helical scans and prevent such acquisitions during axial (nonhelical)scans.
Definition of Pitch RevisitedAs originally defined for SSCT, helical pitch was calculated as table movement per rotation divided by slice thickness. Forexample, with a slice thickness of 5 mm and a table movement of 7.5 mm per rotation, pitch would be 1.5. Because slice thicknessand x-ray beam width are equivalent in SSCT, the value for pitch conveyed important information about the x-ray beam; a pitchof 1.0 meant that the x-ray beams from adjacent rotations were essentially contiguous. Pitches of greater than 1 implied gapsbetween the x-ray beams from adjacent rotations. Pitches of less than 1 implied x-ray beam overlap (and thus double irradiationof some tissue) and so were not clinically used.Applying this definition to MSCT creates confusion and tends to obscure important information. For example, a 4-slice MSCThelical scan with 15 mm of table movement per rotation and a 20-mm-wide x-ray beam (to acquire four 5-mm slices) would yieldthe following pitch calculation based on the earlier definition: pitch = table movement per rotation/slice thickness = 15mm/5 mm = 3.0.
This calculation does not immediately convey the fact that although the pitch is much greater than 1, thereis clearly x-ray beam overlap, because the total width of the x-ray beam is 20 mm and the table is moving only 15 mm per rotation.To address this situation, a new definition of pitch was adopted. In this definition, the denominator is replaced with thetotal thickness of all of the simultaneously acquired slices; that is, if n slices each of slice thickness T are acquired, then the total width is n × T, and the new pitch definition is as follows: pitch = table movement per rotation/( n × T) (beam pitch). Because the original definition is still occasionally referenced, the new pitch definition in the latterequation is called “beam pitch.” The original definition is now referred to as “detector pitch,” on the basis of the ideathat slice thickness (in the denominator of the original definition) in MSCT is determined by detector configuration. Withthe new definition, beam pitch for the example just given would be calculated as follows: pitch = table movement per rotation/( n × T) = 15 mm/(4 × 5 mm) = 0.75.
Because beam pitch conveys similar information for MSCT as the original definition did forSSCT, it is the preferred definition in most situations ( ). Pitch and z Sampling in Helical MSCTClinical pitch selection in SSCT was generally straightforward. Typically, only 2 pitches were commonly used: pitch = 1 forbest quality and pitch = 1.5 when more z-axis coverage was needed in a shorter time (because of either total scan time or x-ray tube heating constraints).
Pitchesof less than 1 were not used. In contrast, commonly used beam pitches in MSCT may seem odd (e.g., 0.9375, 1.125, or 1.375)and are very often less than 1.
Before helical pitch in MSCT is discussed, the basic trade-off involving pitch selection isreviewed (see the first article in this series ( ) for a more complete discussion).Because of continuous table movement, no specific slice position along the z-axis actually contains sufficient data (i.e., transmission measurements along ray paths through the slice at sufficient locationsand angles) to reconstruct an image. Rather, required measurements are estimated by interpolation between the nearest measurementsabove and below the slice plane that are at the same relative position and angle. The distance along the z-axis between these measurements that is available for interpolation is referred to here as the z-spacing. Interpolated data may be inaccurate if anatomy changes significantly within the z-spacing, leading to streak or shading artifacts.
Helical interpolation artifacts often appear as (and are referred to as)“windmill” artifacts, because when the helical images are paged through quickly, the streak or shading artifacts seem to rotatelike the vanes of a windmill. The likelihood and severity of helical artifacts increase with increasing z-spacing, because anatomy is more likely to change abruptly over distance. Increasing pitch (to reduce either scan time orx-ray tube heating) increases distance between interpolated measurements, so that the likelihood of helical artifacts increases.In helical SSCT, slice data are interpolated between equivalent rays separated by a full rotation (360° apart) or betweenparallel opposed rays (180° apart). These 2 interpolation schemes are referred to as 360° linear interpolation (360° LI) and180° linear interpolation (180° LI), respectively. Because parallel opposed rays in SSCT interleave those separated by 360°,z-spacing for 180° LI averages half that for 360° LI ( provides clarification of this statement). Because of its smaller z-spacing and therefore reduced helical artifacts, 180° LI is generally preferred over 360° LI.
Z-Spacing in helical SSCT is minimized by use of 180° LI and a pitch of 1 (assuming that pitches of less than 1 are avoided),in which case the average z-spacing equals d/2, where d is the slice thickness ( ).In MSCT, 180° LI and a pitch of 1 does not improve z-spacing, as demonstrated by the 4-slice example in. With detector collimation represented as “d,” the detectors move relative to the patient by 4d (4 slice thicknesses) inone full 360° rotation. After a half rotation (180°), the detectors move 2d. Unlike the situation in SSCT, 180° opposed raysnow duplicate, rather than interleave, those separated by a full rotation, resulting in a z-spacing equal to d. Suppose, instead, that a value of 3d (3 slice thicknesses) is used ; then, 180° opposed rays interleave those 360° apart and provide a z-spacing of d/2, equivalent to that achieved in SSCT with a pitch of 1 and the same slice thickness. A “rule” for this particularscheme is to overlap one detector width (one slice thickness). For the 4-slice example, this corresponds to a pitch of 3/4= 0.75; for a 16-slice scanner, the pitch is 15/16 = 0.9375; and for a 64-slice scanner, the pitch is 63/64 = 0.9844.
Notethat such pitches move closer to 1 (smaller fraction of beam overlap) as the number of simultaneously acquired slices increases(more information regarding this point is provided later in this article) ( ). FIGURE 7.z-Spacing in helical CT. (A) Minimum z-spacing equal to d/2 (d = slice thickness) is achieved in SSCT with pitch of 1 and interpolation between interleaved parallelopposed rays. (B) With pitch of 1 in MSCT, parallel opposed rays overlap rather than interleave, giving z-spacing equal to d.
(C) z-Spacing equivalent to that in SSCT is achieved with pitch that overlaps one slice thickness but results in double irradiationof some tissue. Reduced z-spacing can also be achieved with other pitches. Det = detector; rot = rotation. Selection of Helical MSCT Pitch and Data InterpolationAlthough illustrating why MSCT helical pitches may seem odd or may be less than 1, the “one-detector overlap” method is notnecessarily optimal or preferred. Some MSCT scanner manufacturers recommend alternative pitch strategies, whereas still othersmaintain that with appropriate interpolation procedures, all (reasonable) pitches are equally good (,). The helical scans shown in help clarify these differences and highlight an important distinction between SSCT and MSCT data interpolation. If a 4-slicescanner uses a detector configuration of 4 × 5 mm to acquire four 5-mm-thick slices , then slice interpolation proceeds as described earlier (usually with 180° LI), preferably with a pitch providing a smallerz-spacing (such as the one-detector overlap mentioned earlier) to reduce interpolation artifacts.
In, with a detector collimation of 4 × 1.25 mm, 13 detector samples lie completely or partially within the 5-mm slice planeto reconstruct (depending on the slice position relative to the overlapping detectors, 11–13 samples may lie within the sliceplane). It is clear that this example no longer involves a simple interpolation between the 2 nearest points but rather requiresan appropriate combination of all measurements lying within the slice.
With appropriate weighting, these measurements maybe combined to form a 5-mm-thick slice measurement with a well-shaped beam profile. This process is referred to as “ z filtering.” With a pitch of 1 rather than 0.75, z-spacing increases from 0.625 mm (i.e., d/2) to 1.25 mm (i.e., d) and leads to only minor degradation of the slice profileshape and increased helical artifacts. For 16-slice scanners (and more), the detector collimation during helical scans is1.5 mm or less (depending on the model and beam width), with correspondingly small differences in the slice profile shapeversus pitch. Helical MSCT Pitch and mAsLow contrast sensitivity (the ability to resolve low-contrast structures) depends on CT image noise. CT image noise originatesprimarily from quantum mottling, which depends on the number of x-ray photons contributing to the image (the appearance ofimage noise also depends on the sharpness of the reconstruction filter used). To understand how various factors affect CTimage noise, it is easiest to consider how many x-ray photons contribute to each detector measurement.
Relevant factors includekVp (with higher kVp, more x-rays penetrate the patient to reach the detectors), slice thickness (the detectors collect morephotons over thicker slices), x-ray tube mA (a higher x-ray intensity increases the number of detected x-rays proportionally),and rotation time (faster rotation corresponds to shorter detector sampling times). The last 2 factors are often taken togetheras mAs (see the second article in this series ( ) for a more complete discussion).Helical SSCT slices are reconstructed from data interpolated between the 2 nearest parallel ray measurements (usually with180° LI). Therefore, the number of x-ray photons contributing to each interpolated sample (and therefore to reducing imagenoise) is a linear combination of 2 detector measurements, regardless of pitch. That is, helical SSCT image noise is unaffectedby pitch ( ). (Note, however, that the interpolation algorithm does affect image noise; fewer rays contribute to images when 180° LIis used than when 360° LI is used, so that 180° LI images are somewhat noisier). Pitch does affect image noise in helicalMSCT if slice measurements are formed from many detector samples.
For example, the 5-mm slice in is formed from a combination of 11–13 detector measurements. If the average x-ray flux reaching each detector element isn, then the number of x-ray photons contributing to the calculated ( z-filtered) sample is between 11 n and 13 n. In comparison, a pitch of 1.5 and a detector collimation of 4 × 1.25 mm results in only 5–7 detector measurements lyingwithin the 5-mm slice plane and thus contributing to each slice sample.
For the same average detector flux n as that used in the earlier example, the number of contributing x-ray photons is 5 n–7 n. That is, fewer x-ray photons contribute to each calculated slice sample for larger pitches, leading to noisier images.In general, the number of photons contributing to images decreases linearly with helical MSCT pitch if the same x-ray techniquesettings (kVp and mAs) are used. As discussed later in this article, with the same x-ray technique factors, patient radiationdose (CT volume dose index CTDI vol) also decreases linearly with pitch (in effect, the same amount of energy is spread over more tissue in the z-direction). Therefore, a practice adopted by some manufacturers is to specify “effective”mAs(mAs eff), calculated as rather than actual mAs, during examination prescription. MAs eff is chosen to maintain the same level of image noise regardless of selected pitch. For example, with 1-s rotation times, amAs eff of 240 uses 240 mA with a pitch of 1 (mAs eff = 240/1 = 240) but uses 300 mA with a pitch of 1.25 (mAs eff = 300/1.25) and 200 mA with a pitch of 0.83 (mAs eff = 200/0.83). MSCT RADIATION DOSIMETRYAlthough axial and helical MSCT involves a more complex data collection process, measuring and specifying patient radiationdoses in MSCT are no different from in SSCT.
For both axial and helical dosimetry purposes, detector collimation is ignoredand an MSCT scanner is treated as a single-slice scanner with a slice thickness equal to the full collimated x-ray beam width.For example, an MSCT scan with a detector collimation of 4 × 2.5 mm (total beam width of 10 mm) would be treated for dosimetrypurposes as an SSCT scan with a slice thickness of 10 mm (see the second article in this series ( ) for a complete discussion of CT dosimetry). For axial scans, therefore, the weighted CTDI CTDI w for a detector collimation of 4 × 2.5 mm (the slices from which may be combined to form 10-mm slices) is equivalent in principleto that of a 10-mm SSCT slice. Similarly, the CTDI vol for helical scans is obtained from axial CTDI w measurements at the same beam collimation by dividing the axial CTDI w by the pitch.There are, however, certain factors that reduce the dose efficiency of MSCT relative to SSCT. In addition, certain MSCT practicesthat were uncommon or nonexistent in SSCT may lead to increased patient radiation doses. These factors and issues are describedin the following discussion. MSCT Dose EfficiencyDose efficiency refers to the fraction of x-rays that reach the detectors and that are actually captured and contribute toimage formation.
Dose efficiency has 2 components: geometric efficiency and absorption efficiency. Absorption efficiency refersto the fraction of x-rays that enter active detector areas and that are actually absorbed (captured).
Absorption efficienciesare similar for all SSCT and MSCT scanners that have solid-state detectors. Geometric efficiency refers to the fraction ofx-rays that exit the patient and that enter active detector areas.Two aspects of MSCT reduce its geometric dose efficiency relative to that of SSCT. The first is the obvious necessity fordividers between individual detector elements along the z-axis, which create dead space that did not exist within SSCT detectors in the z-direction (there is, of course, dead space from detector dividers within the slice plane for both SSCT and MSCT). These dividersare visible in as the thin, lighter lines between the small detector elements. Depending on detector design and element size, dead spaceassociated with the dividers can represent up to 20% of the detector surface area. That is, up to 20% of x-rays exiting thepatient will strike dead space and not contribute to image formation. Because these dividers must satisfy anti–cross talkand physical separation requirements, divider width generally remains unchanged as detector elements are made smaller (comparethe dividers between the 0.625-mm elements and the 1.25-mm elements in ).
Therefore, the dividers represent a larger fraction of detector surface area for smaller elements, leading to lower geometricefficiency. Reducing detector element width from 1.25 mm to 0.635 mm or from 1.5 mm to 0.75 mm approximately doubles the amountof dead space. Geometric efficiency loss is fixed by MSCT detector design and cannot be recovered ( ).The second factor that reduces MSCT geometric efficiency is associated with the x-ray beam width. In SSCT, the beam widthis taken to be the z-axis dose profile width measured at the isocenter (i.e., at the axis of rotation) between profile points corresponding to50% of the maximum intensity (referred to as the FWHM). A collimator is designed such that the profile FWHM corresponds tothe desired slice thickness.
A illustrates such a dose profile for a 10-mm-wide beam used to irradiate an MSCT detector collimation of 4 × 2.5 mm (to acquirefour 2.5-mm slices). Four sections of this profile are shaded to emphasize the x-ray flux contributing to each of the 4 slices.Note that the 2 outer slices receive fewer x-rays—and therefore exhibit more quantum mottling—than do the 2 inner slices.This undesirable situation arises because the 2 outer slices are partially irradiated by the dose profile “edges” (which correspondto the beam penumbra). Providing equivalent radiation to all 4 slices requires that the x-ray beam be widened such that all4 slices are irradiated by the inner, nonpenumbra region, as illustrated in. In effect, the penumbra regions that contributed to SSCT images cannot be used in MSCT and must be discarded ( ). The size of the beam penumbra is related to the collimator design and the focal spot size and changes only moderately at differentbeam widths. As a result, the fractional loss of dose efficiency associated with the discarded penumbra becomes smaller forlarger beam widths, because the penumbra represents a smaller fraction of the total x-ray beam width.
A consequence is thatCTDI values in MSCT are higher for smaller beam collimation values. In comparison, CTDI values in SSCT are nearly independent of slice thickness (and thus beam width, as defined earlier),except in certain cases of thin slices (∼1 mm) for which the beam width deviates from the earlier definition.
As beams becomeeven wider for higher-slice-count scanners (currently up to 40 mm for 64-slice scanners), the geometric efficiency loss associatedwith the penumbra becomes less and less of a factor. Other Dose-Related Issues in MSCTAs mentioned earlier, helical pitches of less than 1 are sometimes used in MSCT, leading to double irradiation of some tissue.More generally, because CTDI vol increases linearly as pitch is reduced, mAs—and therefore patient dose—should be reduced for smaller pitches to compensate(as mentioned earlier, some systems do this automatically by specifying mAs eff rather than actual mAs).Other dose-related issues are not specific to MSCT but rather are associated with the rapid growth of CT spawned by MSCT technologyand the new applications that it permits.
This rapid growth has created increasing concern about population exposure fromCT. Rather than a detailed description of these issues, a brief list of some operational practices to help minimize patientradiation dose is provided here; some of these were described in more detail in the second article in this series ( ). Patient Size and Technique Selection.With the same scan technique factors (kVp, mA, rotation time, and slice thickness), patient radiation dose (CTDI) is considerablyhigher for smaller patients (,).
For example, CTDI w measured in the standard 16-cm dosimetry phantom (representing a small pediatric patient) is nearly double that measuredin the standard 32-cm dosimetry phantom (representing a medium to large patient) with the same technique factors. It is importantto adjust mAs appropriately for patients of different sizes, especially pediatric patients. To facilitate this practice, mostmanufacturers now provide weight- or size-based scan techniques for pediatric patients. Automatic Exposure Control (AEC).Another way in which to tailor a technique appropriately to patient size is to use CT AEC, now available on many scanners(and often referred to as auto-mA).
In a fashion analogous to radiographic AEC, CT AEC automatically selects scan mA on thebasis of patient attenuation estimated from scout views. This process automatically provides an appropriate technique notonly for each patient but also for each individual slice (or each individual rotation during helical scans); for example,higher mA will be used through the diaphragm and abdomen than through slices containing more lung tissues. In clinical practice,the mA automatic selection process is usually based on a user-specified acceptable image noise level. Rotational AEC.Some systems now allow mA adjustment not only for each slice or rotation but also for individual views (angles) during a singlerotation. This feature is most useful for anatomic regions that are far from “round,” such as the hips or shoulders.
In thesecases, image noise (and image quality) is dominated by the very low x-ray intensities transmitted though the lateral viewsof the patient. Conversely, patient dose tends to be dominated by the higher intensities penetrating through the patient fromthe anterior and posterior views. By increasing mA for the “thick” patient views to reduce noise and reducing mA for the “thin”patient views to reduce radiation dose, one may achieve equivalent image quality (relative to nonrotational AEC) with as muchas a 50% dose reduction ( ).
MSCT Beam Width and Radiation Dose.Because radiation doses are higher for thinner beam widths in MSCT , thinner beam widths should be avoided when possible. For example, for a 4-slice scanner, the use of a detector collimationof 4 × 1 mm or 4 × 1.25 mm (total beam width of 4–5 mm) should be avoided unless off-axis reconstructions are planned (forwhich the thinner slices are superior). Similarly, for a 16-slice scanner, a detector collimation of 16 × 1 mm, 16 × 1.25mm, or 16 × 1.5 mm will yield lower doses than a detector collimation of 16 × 0.5 mm, 16 × 0.625 mm, or 16 × 0.75 mm for bothaxial and helical MSCT scans ( ). Advantages of MSCTThe rapid adoption of MSCT technology testifies to its advantages over SSCT. The principal basis of its advantages can bestated as follows: MSCT allows large anatomic ranges to be scanned while simultaneously producing both thin and thick slices.The availability of thick (4–5 mm) slices is important because they are generally preferred for primary interpretation asa result of their lower image noise. Acquiring thin slices is important for 2 reasons: they reduce or eliminate partial-volumestreaks, and they allow for the production of high-quality off-axis (sagittal, coronal, or oblique) or 3-dimensional reconstructions,often with a spatial resolution equivalent to that within the plane of the slice (referred to as isotropic resolution).Although SSCT could, in principle, acquire thin slices, total scan times and subsequent x-ray tube heating prevented practicalthin-slice scanning with thicknesses of less than 2.5–3 mm, except for very limited anatomic ranges. Typical off-axis (coronal)reconstructions available for SSCT and for MSCT are compared in.
The quality of MSCT off-axis reconstructions is such that in some cases, CT image interpretation is beginning to migratetoward primary interpretation from off-axis rather than axial images. Data IssuesThe growth of MSCT has been accompanied by rapid growth in the numbers of images per examination.
The large numbers of imagesassociated with MSCT examinations render film-based interpretation impractical and essentially mandate the use of PACS forefficient and timely examination interpretation. Although it is certainly true that MSCT examination sizes (in terms of imagecounts) are much larger than those for SSCT, significant misunderstandings exist regarding the continued expansion of dataas MSCT continues to evolve.With the introduction of 4-slice MSCT, the sizes of many (but not all) studies increased by 4- or 5-fold or more. This increasewas attributable not so much to collecting 4 slices at once but rather to MSCT scanning of large sections of anatomy withthin slices (2.5 mm or less), relative to SSCT scanning of the same anatomy with 5- to 10-mm slices.
For example, coveringa 40-cm scan range with contiguous 5-mm slices (whether acquired axially or helically) would generate 80 images. MSCT scanningof the same range with a collimation of 4 × 1.25 mm to produce both 1.25- and 5-mm slices would generate 400 images: 320 imageswith a thickness of 1.25 mm and 80 images with a thickness of 5 mm.The reason for the large increase in image counts in the example just given is coverage of the same anatomic range with boththin and thick slices rather than simultaneous acquisition of multiple images. A 16-slice scanner covering the same anatomywith a collimation of 16 × 1.25 mm (again reconstructed into both 1.25- and 5 mm slices) would also yield 400 images—but itwould obtain them 4 times faster (assuming an equivalent rotation time). On the other hand, if the same 40-cm area were scannedwith a collimation of 16 × 0.625 mm to produce 0.625- and 5 mm slices, then 720 images would result (640 images of 0.625-mmslices and 80 images of 5-mm slices). A 64-slice machine scanning the same area to produce 0.625- and 5-mm slices would alsogenerate 720 images. In general, the numbers of images produced by MSCT are associated with detector collimation and how thedata are used rather than with the numbers of simultaneously acquired slices. How many slices an MSCT scanner can acquiresimultaneously affects how fast images are acquired—not how many images are acquired.Although scanners capable of collecting more than 64 slices are now either available or in clinical trials, it is unlikelythat detector element size and therefore minimum detector collimation will continue to shrink.
Because current element sizes(typically 0.5–0.75 mm) are already comparable to detector element sizes within the slice plane (see, for example, the in-planeand z-direction dimensions of the 0.625-mm elements in ), still-smaller z-axis elements lead to diminishing returns. Thus, except perhaps for new and special applications, such as cardiac CT angiography(CTA) (see next section), the numbers of images in MSCT examinations are unlikely to increase significantly. Cardiac MSCTThe application that currently seems to be driving state-of-the-art MSCT is CTA.
Siemens Ct Scan 32 Slice 1
Except for cardiac screening applicationswith electron-beam CT, cardiac imaging has been a difficult hurdle for CT because of its demanding performance requirements.In most cases, the optimal time for scan data collection is during a relatively motionless-heart window of time (lasting about175 ms for a heart beating at 60 beats per minute) occurring at approximately 65%–75% of the R-R interval. Because of thisvery short time interval, CTA examinations are electrocardiographically gated helical scans with data acquired during thesame heart phase over several heartbeats. Small beam pitches (0.25 or less) are used to ensure the collection of data forslice interpolation during appropriate rotational tube positions. For optimal examination quality, data associated with thereconstruction of individual slices should be collected during a single beat, with the entire heart being covered in as fewbeats as possible ( ).Steps toward meeting these requirements have led to faster rotation times and larger z-axis fields of view. Normally, individual cardiac axial images are reconstructed from “half-scan” data (i.e., data collectedover a half rotation 180° plus the fanbeam angle, or about 210°) rather than from 360° of data. Currently, the fastest rotationtimes are on the order of 1/3 s. At this speed, half-scan data can often be acquired during a single beat or, at most, 2 beats.Further speed improvements are difficult to achieve because of both mechanical stresses and x-ray output limitations.
A possiblesolution now in clinical use by one manufacturer involves 2 x-ray tubes and detector arrays, so that each tube–detector pairneed only complete a quarter rotation (taking about 85 ms) to acquire half-scan data.To allow coverage of the entire heart in as few beats as possible, z-axis fields of view have increased to a current maximum of 40 mm. One manufacture has begun clinical trials of a 256-sliceMSCT scanner with a 128-mm z-axis field of view (but with a slower 0.5-s rotation time) ( ). CONCLUSIONMSCT will continue to evolve. Most likely, future MSCT technology advances will be aimed at improving CTA. Except for CTAand a few other examinations that benefit from greater speed or large z-axis fields of view, it is unclear at this point whether additional, significant clinical advantages will accrue beyond thoseof 16-slice MSCT. Because z-axis resolution and scan quality are essentially identical, the gain seems to be mostly in scan times, which are alreadyextraordinarily short for 16-slice scanners. Whether 16 slice, 64 slice, or beyond, MSCT will continue to grow in importanceas a primary diagnostic imaging tool.
Z-Spacing of Parallel Opposed Rays in Helical CTThe spacing along the z-axis of measurements (samples) used to interpolate helical CT slices affects the presence of helical artifacts in images:the smaller the spacing, the less severe the artifacts. Although diagrams such as those in and are often used to illustrate helical z-spacing, they are somewhat misleading. Parallel opposed rays used for interpolation (or z filtering) are separated by a half rotation (180°) only for detectors at the center of the fanbeam detector array (i.e.,at the center of the array within the plane of the slice). Thus, parallel opposed rays occur when the x-ray tube is on theexact opposite side of the patient.For all other detectors, parallel opposed rays occur at some other point during the rotation (either less than or more thana half rotation), depending on the location of the detector within the fanbeam array: the closer the detector is to the endof the array, the farther from a half rotation the parallel opposed ray occurs.
CT Scanners CT Scanner Buyers Guide: Slice Counts and PricingAUTHOR: Jeff Hough January 21, 2019Are you looking to purchase a? If the answer is yes, are you considering a CT scanner with the most advanced technologies available or a machine that delivers the most value to your practice? Whatever your answer may be, it helps to know your options and we are here to help find the right machine to fit your clinical and financial requirements.Determining your clinical requirements is the first, and often most important, step in narrowing down the right scanner. We’re all excited and tempted to purchase the most advanced unit, however the best investment is made by matching the CT to the facility’s requirements. As we highlight below, the price difference between a CT scanner ideal for routine scans may be hundreds of thousands of dollars less than a premium machine designed for specialty procedures. The healthcare experts have broken down the basics you need while consider a new CT scanner.
Slice CountMore slices equal a better CT scanner, right? From a marketing stand point, slice count is a simple metric to compare two scanners.
In practice “better” is subject to utilization and if the scanner produces a quality image for patient care. The term slice count refers to the number of cross-sectional images acquired with each rotation of the CT’s gantry. Advantages of a machine with a higher slice count include: reduced scan times, lower radiation dosages, thinner slices for more detailed images, and capabilities for advanced scans to be performed such as cardiac studies.
The trade-off is a higher price tag and no change in clinical outcome for routine studies. For many facilities, a lower slice count, and a lower price tag, is the best value for the studies they are performing. So How Much Does a CT Scanner Cost?We breakdown CT prices for scanners with 16-slices, 64-slices, 128-slices, and 256-slices. The price ranges are based on new and refurbished existing scanners across the major manufacturers including: GE, Siemens, Toshiba, and Philips. 16-slice CT Scanners:The most common scanner is the 16-slice CT scanner. These units are ideal for general studies and moderate patient volumes.The cost of new and refurbished 16-slice CT scanners are:. New 16-slice costs between $285,000-$360,000.
Refurbished 16-slice costs between $90,000-$205,000Popular 16-slice models include:.64-slice CT Scanners:The 64-slice CT scanner is standard for hospitals, health systems, and imaging centers. Advanced studies, such as cardiac, can be performed due to the reduced scan times and advanced technology. The speed and accuracy that comes with a 64-slice scanner reduces scan times and is suited for practices with moderate to high patient volumes. The cost of new and refurbished 64-slice CT scanners are:. New 64-slice cost between $500,000-$700,000. Refurbished 64-slice cost between $175,000-$390,000Popular 64-slice models include:. GE Revolution.
Siemens Somatom Definitions AS.128+ Slice CT Scanners:Premium scanners with 128, 256, 320+ slices, are most commonly found supporting specialty practices and where patient volumes are high. These scanners are designed to produce crisp images of any organ and generally feature specialty software packages.
The cost of new and refurbished premium CT scanners are:. New 128-slice cost between $675,000-$1,000,000. Refurbished 128-slice cost between $225,000-$650,000. New 256+ slice / dual energy scanners cost between $1,350,000-$2,100,000Popular 128 + slice models include:. GE Revolution CT: ES.
GE Revolution CT: EX. Siemens Somatom Drive. Siemens Somatom Edge. Siemens Somatom ForceSoftware and Hardware Features:CT scanners are easily categorized by slice count but their capabilities are tied to a combination of hardware and software. These features can have a significant impact on the price of a machine. For example, a cardiac software suite can between $35,000 - $100,000, whereas the lung application may result in an extra $15,500-$35,000.
This reinforces the importance of addressing your clinical requirements and the type of studies your facility will be performing before making a purchase.Complementing the detector is the x-ray tube. This is the energy source used to create the image. Depending on manufacturer and model, there may be multiple tube options. These include maximum power output, bearing style, and life expectancy.
Specific to refurbished scanners, tubes with a lower usage generally hold more value than a high usage tube. The price of an x-ray tube can cost between $40,000 - $200,000. Service and Support:The price of a CT scanner is only a portion of total cost of ownership. Maintenance, electricity, site planning, and operation costs are other components.
With a new CT scanner, a 1-year parts & labor warranty is built into the cost. With used and refurbished equipment, service is treated as an additional item. A multi-vendor, OEM service, or Independent Service Provider (ISO) contract can be used. As a benchmark, a full parts & labor service contract through the OEM is typically 10%-14% of the purchase price.
Multi-vendor and ISO contracts can reduce that cost by 20%-30%. Additional savings are possible with parts only, or Time & Material contracts. While there are multiple options available, Meridian Leasing is here to help you make the best choice.This guide should be used as an overview to understanding the different factors that go into the price of a CT machine. The price ranges listed above are averages derived from the current market of the top manufacturers. Your purchase price will depend on the OEM you choose and which scanners are available at the time. Call or send an email to for product and price related details or to speak with a medical equipment specialist.
If you have a specific model in mind, contact us to request a customized quote.