The distribution of radiation dose in CT is clearly different than in radiography, because of the unique way in which radiation dose is deposited. Three aspects of radiation dose in CT that is unique in comparison to x-ray projection imaging. First, because a single CT image is acquired in a highly collimated manner, the volume of tissue that is irradiated by the primary x-ray beam is considerably smaller compared with, for example, the average chest radiograph. Second, the volume of tissue that is irradiated in CT is exposed to the x-ray beam from almost all angles during the rotational acquisition, and this more evenly distributes the radiation dose to the tissues in the beam. In radiography, the tissue irradiated by the entrance beam experiences exponentially more dose than the tissue near the exit surface of the patient. Finally, CT acquisition requires a high SNR to achieve high contrast resolution, and therefore the radiation dose to the slice volume is higher because the techniques used (kV and mAs) are higher. a rough comparison, a typical PA chest radiograph may be acquired with the use of 120 kV and 5 mAs, whereas a thoracic CT image is typically acquired at 120 kV and 200 mAs.
The âComputed Tomography Dose Index (CTDI)â is the essential CT dose descriptor. By making use of this quantity, the first two peculiarities of CT scanning are taken into account: The CTDI (unit: Milligray (mGy)) is derived from the dose distribution along a line which is parallel to the axis of rotation for the scanner (= z-axis) and which is recorded for a single rotation of the x-ray source. Fig.2. illustrates the meaning of this term: CTDI is the comparable of the dose value inside the irradiated slice (beam) that would result if the absorbed radiation dose profile were entirely concentrated to a rectangular profile of width equal to the nominal beam width, with N being the number of independent (i.e. non-overlapping) slices that are acquired simultaneously. Accordingly, all dose contributions from outside the nominal beam width, i.e. the areas under the tails of the dose profile, are added to the area inside the slice it is not necessary to carry out all the scans. Instead, it is adequate to obtain the CTDI from a single scan by acquiring the entire dose profile according to Eqn.1
The CTDI technique uses this long ionization chamber (FIG.1) to integrate the primary and scattered radiation delivered with a single scan (ie, one gantry rotation) and normalizes it to the nominal beam width. This normalization smartly incorporated a scannerâs dose efficiency. That is, if the radiation dose profile from a CT system was unnecessarily wide (ie, the primary beam was wider than the imaged section width FIG.2), the CTDI would be higher than that from a system with a more narrow beam that better matched the width of the imaged. In additional demonstrated that the CTDI could be easily scaled to reflect the common situation when the radiation beams were not contiguous (ie, when there were gaps or overlaps between consecutive rotations of the x-ray tube). Thus, CTDI-based metrics became the reference standard for measuring, comparing, and communicating the radiation output of a CT system. In the early days of CT, direct measurement of the multi-slice average dose (MSAD) was a labor-intensive process. It required multiple scan acquisitions, which placed heavy loads on the x-ray tube. The long scan times necessitated use of dosimeters that could integrate dose accurately over several minutes, such as film or thermoluminescent dosimeters. Conversely, the introduction of CTDI provided a much more practical method with which to estimate the MSAD in FIG.3 and hence compute the radiation output of a CT system. First, although the CTDI could be measured by using only a single rotation of the x-ray tube, it represented the dose from a series of scan acquisitions. Second, it facilitated the use of ionization chambers, making measurements faster and easier to obtain. Because the x-ray beam from a CT scanner was too narrow to completely cover the sensitive volume of existing ionization chambers, a 100-mm-long pencil ionization chamber was developed and the partial irradiation effect corrected on the basis of chamber length and nominal beam width.
5.2 DEFINITION OF DOSE MEASURING QUANTITIES IN CT
Compton scattering is the principal interaction mechanism in CT, so the radiation dose attributable to scattered radiation is sizeable, and it can be higher than the radiation dose from the primary beam. Scattered radiation is not confined to the collimated beam profile as primary x-rays are, and therefore the acquisition of a CT slice delivers a considerable dose from scatter to adjacent tissues, outside the primary beam. Furthermore, most CT protocols call for the acquisition of a series of near-contiguous CT slices over the tissue volume under examination. An example: a protocol in which ten 10-mm CT slices are acquired in the abdomen. The tissue in slice 5 will receive both primary and scattered radiation from its acquisition, but it will also receive the scattered radiation dose from slices 4 and 6, and to a lesser extent from slices 3 and 7, and so on. The multiple scan average dose (MSAD) is the standard for determining radiation dose in CT. The MSAD is the dose to tissue that includes the dose attributable to scattered radiation emanating from all adjacent slices. The MSAD is defined as the average dose, at a particular depth from the surface, resulting from a large series of CT slices. The MSAD could be measured directly by insertion a small exposure meter at a point in a dose phantom, taking a large series of CT scans of the phantom with the meter in the middle of the slices, and adding the doses from all slices. An estimate of the MSAD can be accomplished with a single scan by measuring the CT dose index (CTDI). It can be shown that that the CTDI provides a good approximation to the MSAD when the slices are contiguous. The CTDI measurement protocol seeks to measure the scattered radiation dose from adjacent CT slices in a practical manner. The CTDI is defined by the U.S. Food and Drug Agency (AAPM REPORT NO 96) as the radiation dose to any point in the patient including the scattered radiation contribution from 7 CT slices in both directions, for a total of 14 slices, this definition is referred to as the CTDlFDA, Medical physicists usually measure the CTDI with the use of a long (l00 mm), thin pencil ionization chamber. The pencil chamber is long enough to span the width of 14 contiguous 7-mm CT scans and provides a good estimation of the CTDI FDA for 7- and 8-mm slices. A single CT image is acquired at the center of the pencil chamber, and the CTDI is determined from this single CT scan. To calculate the CTDI, all of the energy deposition along the length of the ion chamber is assigned to the thickness of the CT slice. Radiologists and technologists should be aware and familiar with the dose indices normally displayed on the CT scanner console. These indices include the volumetric CT dose index (CTDIvol) and the dose-length product (DLP).
The CTDIvol, which was introduced to take into account the pitch of helical acquisitions, represents the average dose delivered within the reconstructed section, and is, calculated as the weighted CTDI divided by the pitch. The DLP is the CTDIvol multiplied by the scan length. It gives an indication of the energy imparted to organs, and can be used to assess overall radiation burden associated with a CT study. CT scanners now routinely record the CTDIvol, and, in some cases, the DLP. Although the CTDIvol is not the dose to a specific patient, it is an index of the average radiation dose from a CT series. For each protocol selected, and for each patient, the dose indices displayed on the control panel should be carefully monitored and determined to be within a sensible range to prevent accidental overexposure.
In dose measurement the radiation dose attributable to scattered radiation is considerable, and can be higher than the radiation dose from the primary beam. The CTDI measurement protocol seeks to measure the scattered radiation dose from adjacent CT slices in a practical manner. Medical physicist usually measures the CTDI with the use of a long thin pencil ionization chamber, a single CT image is acquired at the center of the pencil chamber, and the CTDI is determined from this single CT scan. To calculate the CTDI, all of the energy deposition along the length of the ion chamber is assigned to the thickness of the CT slice.
The CTDI is mathematically defined as:
CTDI=1/NT â«_(-â)^ââ’ãD(z)dzã 1
D(z) = The radiation dose profile along the Z-axis
N = Number of tomographic sections imaged in a single axial scan which is equal to the
Number of data channels used in a particular scan. The value of N may be less
Than or equal to the maximum number of data channels available on the system,
T = the width of the tomomgraphic section along the Z-axis imaged by one data
Channel, in multi slice CT several detectors element may be group together to form
One data channel, In single slice CT the Z-axis collimator (T) is nominal scan width.
The CTDI is always measured in the axial scan mode for a single rotation of the x-ray source, and theoretically estimates the average dose within the central region of a scan volume consisting of multiple, contiguous CT scans [Multiple Scan Average Dose (MSAD)]. For the case where the scan length is sufficient for the central dose to approach its asymptotic upper limit. The MSAD represents the average dose over a small interval (â’I/2, I/2) about the center of the scan length (z = 0) for a scan interval I, but requires multiple exposures for its direct measurement. The CTDI offered a more suitable yet nominally equivalent method of estimating this value, and required only a single-scan acquisition, which in the early days of CT, saved a considerable amount of time.
FIG.2. Schematic illustrates the profile of radiation dose delivered during a single CT scan. The CTDI equals the shaded area divided by the section thickness (NT). Radiation dose profile along a line perpendicular to the scan plane shows a peak dose level at the center of the primary beam and long dose tails caused by scattered radiation. NT = nominal beam width.
FIG.3. Schematic illustrates the profile of radiation delivered during multiple CT scans. The radiation dose profiles from nine adjacent transverse CT scans along a line perpendicular to the transverse scans, when summed, produces the MSAD profile. The value of MSAD is the average value of this profile over one scan interval in the central portion of the profile.
FIG 5. RELATIVE DOSE OF DIFFERENT PATIENT SIZE
5.2.1 COMPUTED TOMOGRAPHY TISSUE WEIGHTED FACTOR (CTDIW)
The CTDI varies across the field of view (FOV). For example, for body CT imaging, the CTDI is typically a factor or two higher at the surface than at the center of the FOV. The average CTDI across the FOV is estimated by the Weighted CTDI (CTDIw) where:
CTDI_W= 1/3 CTDI_100center+2/3 CTDI_100periphery 2
CTDI_100= 1/NT â«_(-50mm)^50mmâ’ãD(z)dzã 3
The values of 1/3 and 2/3 approximate the relative areas represented by the center and edge
Values, CTDIw is a useful indicator of scanner radiation output for a specific kVp and mAs.
According to IEC 60601-2-44, CTDIw must use CTDI100 as described above and an f-factor for air (0.87 rad/R or 1.0 mGy/mGy).
5.2.2 COMPUTED TOMOGRAPHY VOLUME FACTOR (CTDIVOL)
To represent dose for a specific scan protocol, which almost always involves a series of scans, it is essential to take into account any gaps or overlaps between the x-ray beams from consecutive rotations of the x-ray source. This is accomplished with use of a dose descriptor known as the Volume CTDIw (CTDIvol), where:
CTDI_VOL= (NÃI)/IÃCTDI_W 4
I = the table increment per axial scan (mm) and the pitch in given as:
A large CT pitch factor, i.e. greater than one, implies the presence of interspaces between successive beam rotations and contributes to the ability to scan a long range and to achieve a short scan time and breathe hold. However in some cases image artifacts and deterioration of z-axis resolution might compromise image quality. For a specific clinical problem, any change in the CT pitch factor should be compensated for by a corresponding change in radiographic exposure in order to yield constant volume radiographic exposure. This applies particularly to those CT scanners where the operator selects radiographic exposure instead of the pitch corrected radiographic exposure (volume radiographic exposure). However, any change in CT pitch factor is automatically compensated for in the case of those CT scanners where the operator directly selects volume radiographic exposure.
PITCH= I/(N Ã T) 5
Thus volume CTDIVol in term of pitch is given as:
CTDI_Vol= 1/PITCH ÃCTDI_W 6
CTDIw represents the average absorbed radiation dose over x and y directions at the center of scan from a series of axial scans and CTDIvol represent the average absorbed radiation dose over x,y and z direction. The CTDIvol provides a single CT dose parameter, based on a directly and easily measured quantity, which represents the average dose within the scan volume for a standardized (CTDI) phantom. The SI units are milligray (mGy). CTDIvol is a useful indicator of the dose to a standardized phantom for a specific exam protocol, because it takes into account protocol-specific information such as pitch. Its value may be displayed prospectively on the console of newer CT scanners, though it may be mislabeled on some systems as CTDIw. The IEC consensus agreement on these definitions is used on most modern scanners. While CTDIvol estimates the average radiation dose within the irradiated volume for an object of similar attenuation to the CTDI phantom, it does not represent the average dose for objects of substantially different size, shape, or attenuation or when the 100-mm integration limits omit a
Considerable fraction of the scatter tails, further, it does not indicate the total energy deposited into the scan volume because it is independent of the length of the scan. That is, its value remains unchanged whether the scan coverage is 10 or 100 cm. It estimates the dose for a 100-mm scan length only, even though the actual volume-averaged dose will increase with scan length up to the limiting equilibrium dose value.
3.2.4 DOSE-LENGTH PRODUCT (DLP)
To better represent the overall energy delivered by a given scan protocol, the absorbed dose can be integrated along the scan length to compute the Dose-Length Product (DLP), where:
DLP(mGy-cm)= CTDI_volÃScan Length 7
The DLP reflects the total energy absorbed (and thus the potential biological effect) attributable to the complete scan acquisition. Thus, an abdomen-only CT exam might have the same CTDIvol as an abdomen/pelvis CT exam, but the latter exam would have a greater DLP (as shown in Table 1), proportional to the greater z-extent of the scan volume. The implication of over ranging with regard to the DLP depends on the length of the imaged body region. For helical scans that are short relative to the total beam width, the dose efficiency (with regard to overranging) will decrease. For the same anatomic coverage, it is generally more dose efficient to use a single helical scan than multiple helical scans.
5.3 PATIENT DOSE
For a CT scan of length (L), starting in the mid-thigh region, the normalized absorbed dose to an organ [Dâorgan (L)] is defined as the organ dose divided by the corresponding CTDIvol. Where CTDIvol is the volume used to perform the specific examination. Values of the DLP can be used to estimate patient effective dose. The effective dose (E) is generally recogÂ¬nized as a parameter that is directly related to the patient stochastic risk, and permits direct comparisons to be made for different types of radiation exposure, Effective doses from MSCT can easily be compared with those for other exposure conditions, such as exposure during radiography or fluoroscopy, exposure to radionuclide and exposure to natural sources of radiation. For the calculation of effective dose, the average absorbed doses to the 20 most sensitive organs must be assessed, which is not feasible in clinical practice. In general, assessment of effective dose is based on corresponding measurements of the CTDIvol or CTDIair. Effective dose is then calculated from these CTDI values using established conversion factors.
DLP can be used to estimate patient effecÂ¬tive doses when the radiation used to perform a given CT examination (i.e., CTDIvol and DLP) remains constant, organ and effective doses are reduced with increasing patient size, and vice versa(as in FIG.5). The patient mass in a CT slice may be estimated from the cross-sectional area, and the correÂ¬sponding average Hounsfield Unit (HU) value, assuming that the HU value is proportional to physical density. This latter approximation is reasonable because at the high photon enerÂ¬gies encountered in CT, most interactions are Compton scatter where the probability of an interaction is directly proportional to patient physical density. Modeling of patients (as a mass equivalent cylinders of water) allows relative doses (as a function of water cylinder diameter) to be used to estimate how organ and effective doses vary with patient size. The relative dose is defined as the ratio of the dose in a cylinder of water that is equivalent to a patient of a speciÂ¬fied size divided by the corresponding dose in a cylinder that is equivalent to a standard individÂ¬ual.
FIG.5 shows relative dose (mean patient dose per 1 mGy of scanner output, CTDIvol) for an abdominal CT scan and different patient sizes (here represented by the sum of anterior posterior [A/P] and lateral dimensions). Over the range of patient sizes from a newborn to a large adult, relative dose is exponentially related to patient size. For a patient with an anterior posterior dimension of 30 cm and a lateral dimension of 40 cm, the anterior posterior + lateral value would be 70 cm and the mean patient dose in the center of the scan range would be approximately equivalent to the CTDIvol value reported on the console. For a neonate having both anterior posterior and lateral dimensions of 10 cm, the anterior posterior + lateral value would be 20 cm and the mean patient dose in the center of an abdomen scan would be about 2.3 times the displayed CTDI value, for body CTDIvol measurements made by using a 32-cm phantom. CTDIvol measurements made on the basis of 16-cm phantoms would require a different scale factor.
5.4 DOSE ESTIMATION WITHIN THE PATIENT:
Although the use of CT in medical diagnosis delivers radiation doses to patients that are higher than those from other radiological procedures, lack of optimized protocols could be an additional source of increased dose. The aim of this study is to determine the role of CTDI in estimating the patients in CT Scanners examinations. The patient organ doses can be estimated using measurements of CT dose indexes (CTDI), exposure-related parameter.
A major factor contributing to the difficulty of determining the dose for a specific patient in the variation in distribution both within a slice and also along the length of the body, within the plane of each slice there is variation in the dose because of the shape and size of the body and internal composition. The variation can be different for heads compared to the larger sections of the body. Most of the dose is deposited in the slice that is being scanned but there is also some radiation that scatters out of the slice and produces some dose in the adjacent tissues. This scattered radiation complicated the process of determining the dose when more than one slice is imaged as in the usual procedure. An additional complication is the fact that we cannot place instruments, or dosimeters, directly into the body to make measurements. Because of this combination of factors and challenges the dose values for a specific patient are determined or estimated through a series of external measurements and calculations.
Because it is not practical to measure directly the dose deposited in a patient’s body an indirect, and somewhat complex process is used. We will now go thought that process which takes into account some of the variable factors that must be considered. The first step is to measure the dose in a phantom that represents a patient body. The typical phantom is an acrylic cylinder that has the same x-ray absorption properties as soft tissue. It is described as being a “tissue equivalent” material with approximately the same density and effective atomic number (Z) as soft tissue. A smaller phantom is used to represent a head and a larger one for the abdominal section. A dosimeter, typically an ionization chamber, is placed in the phantom and a one-slice scan is performed and the dose measured.
Now to consider the first problem we observed earlier some radiation scatters out of the scanned slice and exposes the adjacent tissue. This becomes significant in the typical scan which consists of many slices. The question (?) is “what is the dose within one slice” from both the direct x-ray beam through the slice and the scatter from other slices? The problem is we do not have any practical method to measure the dose in any one slice when it is within the midst of other slices that would expose the dosimeter to their direct radiation.
Now for the solution…it has been demonstrated that if a dose measurement is made for just one scanned slice the scattered radiation out of the slice and measured by the dosimeter will give a good approximation to the dose within a slice including the scattered radiation from adjacent multiple slices as shown in FIG.6.
Since this is not an actual direct measurement, but an indirect approximation, under multiple slice conditions it cannot correctly be referred to as “dose”. The value measured by this process and used as an estimate of the actual dose is designated as the:
Computed Tomography Dose Index (CTDI). Therefore, CTDI values, and not actual dose values, are used to describe the radiation to patients.
5.5 CTDI AND THE ESTIMATE OF PATIENT ORGAN DOSE IN CT
It is well known that patient doses from CT procedures are relatively higher than doses from other imaging modalities based on ionizing radiation. For example, one CT examination of the chest delivers about 400 times the dose delivered by a conventional chest X-ray examination. Therefore, although CT represents less 10% of the total number of medical X-ray procedures worldwide, this high dose procedure contributes about 34% of the annual collective dose from all medical X-ray examinations to the population. This contribution is predictable because it results from a combination of high dose per examination and frequent use of CT examination in diagnosis Increased use of this high dose procedure has been of great concern globally because of the high possibility of inducing undesired health effects, such as induction of cancer, in patients. Of prime concern is the significant radiation dose delivered to superficial radiosensitive organs such as the eye lens, breast, and thyroid, which are, unfortunately, irradiated during radiological procedures of the head, chest, and cervical spine. The implication of some of these exposures, for example, to the breast and eyes, is the potential increase in the risk of breast cancer and cataract formation in the population, since radiation exposure of different organs leads to different health effects.
The most useful way to assess organ doses is either by direct measurement (on patients using thermoluminescent dosimeters (TLDs) or on phantom using either an ionization chamber or TLDs) or by indirect measurement through measurement of CT dose indexes (CTDI). The standardization of the CTDI phantoms marked a crucial step in quantifying the radiation output of a CT scanner. This is because the primary beam emitted from the scanner (originally a relatively thin fan beam, which with current technology has expanded to cone beams of up to 16 cm width along the patient longitudinal axis) produces a substantial amount of scattered radiation when it interacts with the patient. Hence, consistent radiation output measurements required consistent phantoms.
Early estimates of dose from a CT examination did not use the CTDI methodology and measured only the dose from a single scan acquisition. Specifically, only the peak radiation dose emitted by the scanner from a single tube rotation and at a single table position was measured, and this underestimated the dose delivered to a typical adult patient by a factor of two to three. The reason for this underestimation was that the measurement neglected the âtailsâ of the dose distribution caused by scattered radiation produced from scans at adjacent table positions (Fig.2). Because most clinical examinations involve multiple scans (ie, gantry rotations) as the patient is translated through the gantry, the dose distribution to the patient is the sum of the overlapped âsingle-scanâ dose distributions (Fig.3). For examinations with a sufficient number of scans, the average dose over the central scan width of the imaged anatomy will reach an equilibrium value, which is referred to as the multiple scan average doses (MSAD) (Fig 3). In the early days of CT, direct measurement of the MSAD was a labor-intensive process. It required multiple scan acquisitions, which placed heavy loads on the x-ray tube. The long scan times necessitated use of dosimeters that could integrate dose accurately over several minutes, such as film or thermo luminescent dosimeters. Conversely, the introduction of CTDI provided a much more practical method with which to estimate the MSAD and hence quantify the radiation output of a CT system. First, although the CTDI could be measured by using only a single rotation of the x-ray tube, it represented the dose from a series of scan acquisitions. Second, it facilitated the use of ionization chambers, making measurements faster and easier to acquire. Because the x-ray beam from a CT scanner was too narrow to completely cover the sensitive volume of existing ionization chambers, a 100-mm-long pencil ionization chamber was developed and the partial irradiation effect corrected on the basis of chamber length and nominal beam width.
The CTDI technique uses this long ionization chamber to integrate the primary and scattered radiation delivered with a single scan (ie, one gantry rotation) and normalizes it to the nominal beam width. This normalization cleverly incorporated a scannerâs dose efficiency. That is, if the radiation dose profile from a CT system was unnecessarily wide (ie, the primary beam was wider than the imaged section width), the CTDI would be higher than that from a system with a more narrow beam that better matched the width of the imaged section. In addition, CTDI could be easily scaled to reflect the common situation when the radiation beams were not contiguous (ie, when there were gaps or overlaps between consecutive rotations of the x-ray tube). Thus, CTDI-based metrics became the reference standard for measuring, comparing, and communicating the radiation output of a CT system.
In recent years, however, the strengths and weaknesses of the CTDI have been debated. Criticisms of the CTDI are based on two primary arguments: (i) the 100-mm-long pencil ionization chamber used to collect the dose may not be sufficiently long to measure all of the tails of the scattered dose distribution, and (ii) the phantoms used for CTDI measurements are shorter than an adult torso and so do not produce as much scattered radiation as would occur in a typical adult. This means that the average dose (eg, MSAD) that would occur in the much longer âtypical-sizedâ adult torso is underestimated with CTDI measurements in the 14-cm-long body CTDI phantom; the underestimation owing to the use of this phantom can be as much as 40%. Another important limitation of the CTDI concept is that it is not applicable for CT exposures where the patient remains stationary throughout the scan. Whether from wide cone-beam systems that image a large volume without table increment or CT perfusion examinations, the CTDI value presented on the scanner console is an overestimate of both the average dose within the scan volume and the dose to the skin.
These criticisms, however, are based on the belief that CTDI should estimate the patient dose, as opposed to quantifying the radiation output of CT systems. In fact, because patients and the wide range of clinical applications and scan protocols used to scan them vary so dramatically, there is no single phantom that can be used to accurately estimate the dose to all patients. Any dose metric designed to estimate patient dose for a âtypicalâ adult will underestimate the actual absorbed dose for a pediatric patient or overestimate the actual absorbed dose for an obese patient. Instead, because the volume CTDI (CTDIvol) is displayed on the scanner console before the initiation of a scan (to allow the operator to confirm that the proper scanner output is programmed) and recorded as part of the patientâs examination information, many users incorrectly assume that it is the dose to that particular patient. The CTDI values are included in either a screen-captured âpatient dose reportâ or a structured Digital Imaging and Communications in Medicine dose report, which reinforces the incorrect belief that CTDI is a measure of patient dose. In fact, the actual dose to any given patient is directly dependent on the size and shape of the patient. The CTDIvol is a standardized measure of the radiation output of a CT system, measured in a cylindrical acrylic phantom that enables users to gauge the amount of emitted radiation and compare the radiation output between different scan protocols or scanners. Complex calculations are required to map scanner output to patient dose, taking into account the patientâs size, irradiated organs, body composition, and scan range. Although the need to take patient size into account when estimating patient dose has been well established, the widespread misinterpretation of CTDI as a measure of patient dose continues. Equipped with accurate knowledge of scanner output and estimates of patient size (eg, from the CT radiograph), scan region (eg, thorax or abdomen), and scan length, estimates of patient-sizeâ”specific dose may be determined with an accuracy of approximately 10%. Thus, as long as scanner output continues to be measured and reported by using a standardized, highly reproducible, and pragmatic measurement technique, such as the CTDIvol method, patient dose can be accurately estimated. It is imperative, however, that the community be aware that the CTDI is not patient dose.
CTDIvol provides a very useful way to compare the doses delivered by various scan protocols or to achieve a specific level of image quality for a specific size patient. With use of technique charts and diagnostic reference levels, CTDIvol can be used to prescribe the right dose for a specific patient size and diagnostic task. However, CTDIvol cannot be used as a surrogate for patient dose, either in epidemiologic assessments of potential late effects or for potential deterministic effects (eg, skin injury). Neither CTDIvol nor its derivative, dose-length product (DLP, which is the product of CTDIvol and the irradiated scan length), should be used to estimate effective dose or potential cancer risk for any individual patient. The published âk factorsâ used to convert DLP to effective dose all assume a standard-sized patient. The âstandardâ patient used for adult k factors is relatively thin by todayâs standards (nominal body mass of 70 kg). Thus, the patient models used to estimate dose by using DLP do not represent a real patient.
An important implication of the need to take patient size into account, both when estimating patient dose and when prescribing the correct scanner output settings, is that considerable variation in CTDI-based dose metrics can, and should, be expected. Facilities that adjust their CT technique appropriately for patient size, whether with use of manual technique charts or automatic exposure control, will prescribe a wide range of scanner output (CTDI) values. This is a good outcome, reflecting the facilityâs carefulness in âright-sizingâ the dose settings on the basis of specific patient body habitus. in addition, variability in the image quality criteria for various diagnostic tasks and clinical applications introduces inconsistency in the scanner output settings that one should prescribe, even for patients of the same size. For example, scanner output should vary markedly between CT colonography and CT enterography, even for the same patient. Thus, radiation management in CT requires choosing the correct settings for scanner output, not only for patient size but also for the imaging task.
RADIATION DOSE ISSUE WITH MULTI SLICE CT SCANNERS
6.1 MULTI SLICE CT SCANNERS DOSE:
Multi-slice or multi-detector row CT scanners, capable of imaging four simultaneous, parallel slices in a single rotation, were first introduced in 1998. Since then, scanners with 6, 8, 10, 16 20.32, 40, 64, and 128 slice capabilities have become available and have made a marked impact on the role of CT in the diagnostic radiology and hence extended applications of CT. Multi-slice (MS) technology increases the efficacy of CT procedures and offers new promising applications. The expanding use of MSCT, however, may result in an increase in both frequency of procedures and levels of patient exposure. Multi-slice technology has lead to a considerable advance in the capabilities of CT scanners. In terms of essential dose characteristics they are very similar to single slice systems, although some differences exist in terms of z-axis geometric efficiency and detector array geometric efficiency. In helical scanning there may be extra dose from additional rotations at each end of the scan run. In most situations the above factors would increase the dose on a multi-slice scanner by around 20% compared to a single slice system of equivalent design. In some applications they could lead to a doubling or more of dose and in these cases the explanation for that particular technique should be carefully considered.
The amount of radiation dose a patient receives from a CT scan depends upon two key factors, the design of the scanner and also on the way that the scanner is used. The designs of single slice and multi-slice scanners are similar in most aspects that affect radiation dose, but multi-slice scanning can potentially result in higher radiation risk to the patient due to increased capabilities allowing long scan lengths at high tube currents. CT scanners typically display two dose indices, CTDIvol (mGy) and DLP (mGy-cm), along with which two standard CTDI phantoms, 16- or 32-cm diameter, are used to estimate the two indices. CTDIvol is based on radiation dose measurements on an individual scanner completed by a medical physicist with a pencil ionization chamber and either a 16- or 32-cm diameter CTDI cylindrical plastic phantom. CTDIvol represents the radiation dose delivered to a standard plastic phantom from a specific model CT scanner using specified scan parameters: tube voltage, tube current, rotation time, pitch and selected bowtie filter. CTDIvol provides a standardized method to estimate and compare the radiation output of different CT scanners. CTDIvol is defined to be a dose index of CT scanners, not an indication of patient dose. While the measured dose index in the previous example with an ionization chamber, CTDIvol, is very dependent on patient size when scan parameters for all three phantoms are identical, the displayed CTDIvol on a CT scanner is independent of changes in patient size by definition. The displayed scan index assumes the actual scan parameters but by definition assumes the patient is either a 16-cm or 32-cm diameter cylindrical plastic phantom, if the 16 cm is use this overestimates the dose to adult patients and underestimates dose to infants. If the 32-cm phantom is assumed, the scanner displays 18 mGy for all patients; this underestimates the dose to all patients smaller than adults. In either case, the displayed CTDIvol does not represent the patient dose and can mislead the technologist or radiologist.
The increased capabilities of multi-slice scanners, which allow higher mAs values, longer scan lengths and multi-phase contrast studies, have the potential of directly increasing patient doses. Another indirect but significant effect on dose