CT is one of the highest contributor for medical radiation exposure to patients. Some common CT dose descriptors and dose optimizations methods are briefly described in this presentation.

2. Topics
 Radiation Units
 CT Dose Descriptors
 Factors Affecting Dose
 Strategies for reducing dose
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3. Concern
 Because of increasing use of multislice helical CT, CT must be considered a high-dose
procedure. U.S. Public Health Service data suggest that 17% of all x-ray examinations are CT,
yet CT accounts or 50% of total patient effective dose.
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4. CT, 3%
NM, 26%
IF, 3%
CR, 68%
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CT, 49%
NM, 26%
IF, 14%
CR, 11%
Medical exposure of patients, early 1980’s Medical exposure of patients, 2006

5. Concern
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6. RADIATION QUANTITIES AND THEIR UNITS:
Exposure
 Exposure is a measure of the amount of ionization produced in a specific mass of air by x
rays or gamma radiation.
 The unit of exposure is the Roentgen (R), the conventional unit, or the coulomb per
kilogram (C/kg), the SI unit.
 One Roentgen produces 2.58 × 10−4 C/kg of air at standard temperature and pressure.
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7. Absorbed Dose
 The absorbed dose, or radiation dose as it is popularly referred to, is the amount of energy
absorbed per unit mass of material (patient).
 The old unit of absorbed dose is the rad (r), which is equal to an energy absorption of 100 erg/g of
absorber.
 The SI unit of absorbed dose is the Gray (Gy), so named in honor of Louis Harold Gray, a British
radiobiologist who devised ways to measure the absorbed dose.
 One Gray is equal to 1 joule per kilogram (J/kg)
 1 r is equal to 0.01 Gy or 100 r is equal to 1 Gy. 9/21/2021 7

8. Equivalent Dose
 In recognition of the health effects of x-rays, another conversion factor is applied to the absorbed
dose.
 It is the product of the absorbed dose and a radiation weighting factor (wR).
 The unit for the equivalent dose is also the rem or the Sv.
 Equivalent Dose= Absorbed dose x Radiation weighting factor
 100 rem = 1 Sv
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9. Effective dose
 The quantity effective dose, E, is used to quantify the risk from partial-body exposure to that from
an equivalent whole-body dose.
 It is used to take into account the type of radiation and the radio sensitivity of different tissues.
 It extrapolates the risk of partial body exposure to patients from data obtained from whole body
doses to Japanese atomic bomb survivors. This unit is a weighted average of organ doses.
 Old unit of effective dose is the rem, the SI unit is the Sievert (Sv), where 1 Sv = 100 rem.
 The effective dose relates exposure to risk. As noted by Brenner (2006), “…relevant organ doses for
CT examinations are on the order of 15 mSv or less.” Brenner also notes that “…there is good
epidemiologic evidence of increased cancer risk for children exposed to acute doses of 10 mSv (or
more) and for adults exposed to acute doses of 50 mSv (or more).”
 Effective dose = Absorbed dose x Radiation weighting factor x Tissue Weighting Factor
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10. 9/21/2021 10

11. Radiation Weighting Factor (WR)
 Accounts for relative hazard from various forms of radiation
 alpha = 20
 beta = 1
 gamma/x-ray = 1
 neutron = 10
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12. Tissue Weighting Factor
 The tissue weighting factor (WT) is a relative measure of the risk of stochastic effects that might result
from irradiation of that specific tissue. It accounts for the variable radio sensitivities of organs and
tissues in the body to ionizing radiation.
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2007

13. Tissue Weighting Factor
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14. Dose Geometry
 In conventional radiography, the skin of the entrance plane receives 100% of the radiation, and
the percentage falls rapidly as the x-ray beam is attenuated by the patient’s tissue.
 In conventional radiography, it is common for the exit dose to be only 1% (or 1/100) that of the
entrance dose.
 In CT, the difference between the dose at the center and the dose at the periphery is not nearly
as great as that of conventional radiography.
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15. Dose Geometry
 Dose is more uniform in CT than in general radiography for two reasons.
 The beam is heavily filtered as it exits the x-ray tube. This means that fewer low-energy photons
remain. Because the beam is “harder” than that used in conventional radiography, a lower
percentage of the beam will be absorbed or scattered as it passes through patient tissue.
 CT exposure comes from all directions, creating a more uniform exposure.
 The uniformity of the dose decreases as the scan field of view and patient thickness increase.
Therefore, body scans are less uniform than head scans.
 The central dose for a body scan is approximately one-third to one-half that of the peripheral dose
however, organ doses are clearly higher for children compared with (larger) adults.
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16. 9/21/2021 16

17. 9/21/2021 17

18. Comparison of Dose From CT With Dose
From Conventional Radiographic Studies
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19. Z Axis Variations
 In addition to the variations within the scan plane, variations along the length, or z axis, of the
patient, are described by the z axis dose distribution (or radiation profile)
 In general, the tails will contribute approximately 25% to 40% additional dose to the entire study.
Therefore, if a single slice of the chest delivered a dose of 4 cGy, the entire dose from 30 contiguous
slices would range from approximately 5 cGy (i.e., 4 cGy + 1 cGy) to 5.6 (i.e., 4 cGy + 1.6 cGy). As
compared with the dose from a single slice, the overall radiation dose will be higher when multiple
scans are performed.
 So, total dose is the central slice radiation dose, plus the scatter overlap (or tails). This is called the
multiple scan average dose (MSAD).
 Another type of radiation dose measurement in CT is the computed tomography dose index (CTDI).
This allows an estimate of the MSAD to be accomplished with a single scan.
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20. Z Axis Variations
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21. Phantoms for CT Dose Measurement
 There are two standard polymethyl methacrylate (PMMA) dosimetry phantoms; the body phantom is
32 cm in diameter and 15 cm long, and the head phantom is 16 cm in diameter and 15 cm long.
 The head phantom also serves as a pediatric torso phantom.
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22. Phantoms for CT Dose Measurement
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The computed tomography dose index (CTDI) is measured using either a 16-cm or 32-cm
diameter polymethyl methacrylate (PMMA) phantom.

23. Phantoms for CT Dose Measurement
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24. Phantoms for CT Dose Measurement
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25. Phantoms for CT Dose Measurement
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26. CT Dose Descriptors
1. MSAD
2. CTDI100
3. CTDIw
4. CTDIvol
5. DLP
6. SSDE
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27. Multiple Scan Average Dose (MSAD)
 The average cumulative dose to the central slices from a series of slices with constant spacing is
referred to as the multiple-slice average dose.
 The MSAD represents the dose to a specific section location resulting from the scan at that location as
well as contributions from adjacent scan locations since the penumbra from adjacent sections may
contribute to the dose received by the section of interest.
 The MSAD may be 1.25 – 1.4 times (25 to 40% additional dose to the entire study) the single slice dose
depending upon scanner geometry, collimator design, slice spacing and the position of the slice within
the series of slices 9/21/2021 27

28. MSAD
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A series of seven scans spaced (bed indexing) apart along
the z-axis gives seven bell shaped dose distribution
curves (top). When the doses from this series are
summed, the resultant total dose appears as the bottom
curve. The total dose curve has peaks where the bell-
shaped curves overlap. The dotted line through the total
dose curve is the multiple scan average dose.

29. Computed Tomography Dose Index
CTDI100
 The CTDI was originally designed as an index, not as a direct dosimetry method for patient dose
assessment.
 CT Dose Index is the dose a slice gets from the complete series of contiguous slices
 The CTDI is what manufacturers report to the U.S. Food and Drug Administration (FDA) and
prospective customers regarding the doses typically delivered for their machines.
 The basic CTDI100 measurement involves the use of a 100-mm-long cylindrical (“pencil”) chamber,
approximately 9 mm in diameter, inserted into either the center or a peripheral hole of a PMMA
phantom.
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30. CTDI100
 The CTDI100 describes the measurement of the dose distribution, D(z), along the z-axis, from a single
circular rotation of the scanner with a nominal x-ray beam width of nT, where N is the number of
simultaneously acquired slices of thickness T, where the primary and scattered radiation are measured
over a 100-mm length, and where center of the x-ray beam is at z = 0.
 The 100-mm-long thin cylindrical device is long enough to span the width of 14 contiguous 7-mm CT
slices. This provides a better estimate of MSAD for thin slices than that of the single-slice method.
When this method is used it is referred to as the CTDI100.
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31. 9/21/2021 31
Dose that scanning of slice 1 gives to all slices
=
Dose that slice 1 gets from scanning of all slices
(A) Dose to slice 1 from scanning of slice 1. (B)
Dose to slice 1 from scanning of slice 2 equals
dose to slice 2 from scanning of slice 1. (C) Dose
to slice 1 from scanning of slice 3 equals dose to
slice 3 from scanning of slice 1. (D) Dose to slice
1 from scanning of all slices equals dose to all
slices from scanning of slice 1, or total area under
dose profile, which is measured using long
ionization chamber.

32. CTDIw
 The dose for body scans are not uniform across the scan field of view—the dose at the periphery of
the slice is higher than the central dose. The CTDIW adjusts for this by providing a weighted average
of measurements at center and the peripheral slice locations (i.e., the x and y dimensions of the
slice).
 CTDI100 measurements are made for both the center CTDI100,center) and periphery
(CTDI100,periphery). Combining the center and peripheral measurements using a 1/3 and 2/3
weighting scheme provides a good estimate of the average dose to the phantom, giving rise to the
weighted CTDI, CTDI
 CTDIW = (1/3) (CTDI100)center+ (2/3) (CTDI100)periphery
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33. CTDIVOL
 The CTDIvol radiation dose parameter takes the process a step further by taking account the exposure
variation in the z direction.
 In helical CT scanning, the CT dose is inversely related to the helical pitch used, that is, dose ∝
1/pitch
 Because of this relationship, the CTDIW is converted to the volume CTDI (CTDIVOL) using,
CTDIVOL = CTDIW/pitch
 For helical sequences the CTDIVOL is now the preferred expression of radiation dose in CT dosimetry.
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34. CTDIVOL
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35. Limitations of CTDIVOL
 CTDIvol is a dose index and it’s concepts were meant to enable medical physicists to compare the
output between different CT scanners and were not originally intended to provide patient-specific
dosimetry information.
 The dose received by a patient from a CT scan is dependent on both patient size and scanner radiation
output.
 CTDIVOL provides information regarding only the scanner output. It does not address patient size, and
hence does not estimate patient dose.
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36. Limitations of CTDIVOL
 The 32-cm-diameter PMMA body phantom corresponds to a person with a 119 cm (47”) waistline—a
large individual, indeed.
 For smaller patients, the actual doses are larger than the CTDIvol for the same technique factors—thus,
the CTDIvol tends to underestimate dose to most patients.
 For this limitation, researchers have shown that patient size conversion factors can be used with the
CTDIVOL reported on a given CT scanner to produce size specific dose estimates (SSDEs).
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37. Limitations of CTDI
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Both patients scanned with the same CTDIvol
Patient dose will be higher for the smaller patient
CTDIvol = 20 mGy CTDIvol = 20 mGy
120 kVp at 200 mAs 120 kVp at 200 mAs
32 cm
Phantom
32 cm
Phantom

38. Limitations of CTDI
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Smaller patient scanned with a lower CTDIvol
Patient doses will be approximately equal
CTDIvol = 10 mGy CTDIvol = 20 mGy
120 kVp at 100 mAs 120 kVp at 200 mAs
32 cm
Phantom
32 cm
Phantom

39. Size Specific Dose Estimate (SSDE)
 AAPM report 204, 2011 describes a method to calculate SSDE using CTDIvol
 Conversion factors based on patient size (e.g., AP or lateral width, effective diameter) are provided to
estimate patient dose for a patient of that size
 However, SSDE is still not the exact patient dose, as factors such as scan length and patient
composition may differ from the assumptions used to calculate SSDE
 SSDE is not dose to any specific organ, but rather the mean dose in the center of the scanned volume
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40. Size Specific Dose Estimate (SSDE)
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This figure shows data taken from AAPM
Report 204. This curve shows the conversion
factors as a function of the effective diameter
of the patient, relative to the 32-cm-diameter
CTDIvol phantom. The data from this curve
can be used to adjust the displayed CTDIvol
when the effective diameter of the patient is
known.

41. Size Specific Dose Estimate (SSDE)
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42. Size Specific Dose Estimate (SSDE)
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AAPM report 204, 2011

43. Size Specific Dose Estimate (SSDE)
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32 cm
Phantom
32 cm
Phantom
Patients have equivalent SSDE
CTDIvol = 10 mGy
SSDE = 13.2 mGy
CTDIvol = 20 mGy
SSDE = 13.2 mGy
120 kVp at 100 mAs 120 kVp at 200 mAs

44. Why Use CTDIvol?
 CTDIvol is a useful index to track across patients and protocols for quality assurance purposes
 CTDIvol provides information about the amount of radiation used to perform the study
 CTDIvol can be used as a metric to compare protocols across different practices and scanners when
related variables, such as resultant image quality, are also taken in account
 The ACR Dose Index Registry (DIR) allows comparison across institutions of CTDIvol for similar exam
types (e.g., routine head exam)
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45. Dose length product (DLP)
 CTDI in any form is an estimate of average radiation dose only in the irradiated volume for single
slice, so another indicator called Dose length product (DLP) is in use.
 The product of CTDIVOL and the length of CT scan along the z-axis of the patient, L, is the dose
length product (DLP)
 DLP = CTDIVol x L

46. Estimates of Effective Dose from DLP
 It has been shown that the DLP is
approximately proportional to
effective dose (E).
 The slope of the E versus DLP
relationship is called the k value, and
there are different k values for
various CT examinations, as given in
Table 11-5
 Effective Dose= DLP x K
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47. 9/21/2021 47

48. Factors Affecting Dose
1. Radiation Beam Geometry
2. Filtration
3. Detector Efficiency
4. Slice Width and Spacing
5. Pitch
6. Scan Field Diameter
7. Exposure Technique
8. Patient Size and Body Part Thickness
9. Repeat Scans
10. Collimation
11. Localization Scans
12. Number of Detectors
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49. Radiation Beam Geometry
 Theoretically, a rotation arc of only 180° is all that is required to satisfy most construction
algorithms. Most scanners use a 360° tube arc to compensate for radiation beam divergence and
patient motion.
 The extra scanning information improves image quality but increases radiation dose.
Additionally, overscanning, which is the process of using more than a 360° tube arc, is
sometimes used
 Overscanning is also necessary for interpolation of projections in MDCT. Overscans will increase
the radiation dose.
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50. Filtration
 Filtration affects the radiation dose by removing some of the low-energy x-rays.
 These low-energy x-rays are quickly absorbed by the patient.
 Photons are needed to penetrate the patient and strike the detectors, but some of them must be
absorbed to produce radiographic contrast.
 Adding metal filters to the beam permits selective removal of x-rays with low energy and reduces
the radiation dose while maintaining contrast at an acceptable level.
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51. Detector Efficiency
 Detector absorption efficiency affects radiation dose to the patient.
 Less-efficient detectors will require a higher radiation exposure to produce an adequate image.
 Solid state detectors are 90% to 100% efficient, whereas the xenon gas detectors used in older
model scanners are significantly less efficient (60% to 87%) at absorbing x-rays.
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52. Slice Width and Spacing
 Decreasing slice thickness and using contiguous slices will increase the dose because of the increased
amount of scatter radiation to adjacent slices.
 Also, to maintain image quality at the same level, additional radiation is needed for thinner slices.
 Multiple slice examinations using overlapping slices will produce a higher overall dose, whereas
gapped slices will produce a lower overall dose.
 To have the same x-ray intensity reach all of the detectors in an MDCT system, the collimators must
be opened slightly more, allowing the x-ray penumbra to fall outside the active detectors, referred to
as overbeaming
 This contributes to the patient dose but is not used to create an image. The effect is more
pronounced for smaller collimator openings. 9/21/2021 52

53. Slice Width and Spacing
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To have the same x-ray intensity at each of the
detectors in an MDCT scanner, the collimators
must be opened slightly more so that the x-ray
penumbra falls outside the detectors. This
contributes to the radiation dose to the patient
but dose not contribute to the image created.

54. Pitch
 Helical scans performed with a pitch of 1 deliver approximately the same dose as that of
conventional axial CT studies—provided the kVp and mAs values are the same for each mode.
 The pitch has a direct influence on patient radiation dose because as pitch increases, the time
that any one point in space spends in the x-ray beam is decreased.
 It is generally expressed as 0.5, 1.0, 1.5, 2.0
 pitch = 1 - coils of the helix (radiation beam) are in contact
 pitch < 1 - coils of the helix (radiation beam) overlap. Increased patient dose
 pitch > 1 - coils of the helix (radiation beam) are separated. Reduced patient dose
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55. 9/21/2021 55

56. Pitch
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57. Scan Field Diameter
 Holding all technical factors constant, a scan of the head phantom will result in a higher radiation
dose than that of the body phantom.
 The primary difference in results between the head and body phantoms is size.
 The smaller object always absorbs the higher dose and that the difference is at least a factor of two.
 Thus, smaller patients would be expected to absorb much higher amounts of radiation than larger
patients.
 This effect is primarily attributed to the fact that total exposure is made up of both entrance
radiation and exit radiation.
 For smaller patients, the patient has less tissue to attenuate the beam, which results in a much
more uniform dose distribution. Conversely, for a larger patient, the exit radiation is much less
intense as a result of its attenuation through more tissue.
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58. Exposure Technique
 As in general radiography, the technique used to create the CT image affects radiation exposure to
the patient.
 The higher the mAs and kVp settings used to create the image, the higher the dose to the patient.
 The relationship between mAs and dose is linear.
 Assuming that the minimum dose to obtain acceptable image quality has been determined. If this
dose is halved by halving the mAs, a noise increase of 41% can be expected.
 Kvp also affects the radiation dose, although the effect is not linear. With the mAs kept constant,
changing from 120 kVp to 140 kVp increases the radiation dose approximately 30% to 45%.
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59. Effective Milliamperage-Seconds
 The effective mAs is a term used for multislice CT scanners that denotes the mAs per slice. This is
given by the following relationship:
Effective mAs = True mAs/pitch
 This expression simply implies that, to keep the effective mAs constant, as the pitch increases, the
true mAs must be increased as well
 For example, increasing the pitch from, say, 1 to 2, increases the mAs per rotation from 100 to 200
while the relative CTDIvol remains the same
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60. Patient Size and Body Part Thickness
 Large patients or thick body parts require radiographic techniques that increase the radiation
dose to avoid an unacceptable level of image noise.
 In addition, the patient size and body composition may affect the degree of scatter radiation.
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61. Repeat Scans
 Areas of the patient that are rescanned to visualize various stages of intravenous contrast
enhancement or for other technical or clinical reasons receive additional radiation.
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62. Collimation
 Lead collimators are used near the x-ray tube to control the size of the beam striking the patient.
 If the beam is not controlled to match the detector size, there will be additional scatter radiation to
degrade the image; this scenario will result in a higher radiation dose to the patient.
 Collimators may also be used near the detectors for scatter rejection and aperture use.
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63. Localization Scans
 The localization scan performed before scanning, which is often referred to as the scout
image, delivers a very low dose.
 The radiation dose for the scout image is much lower than that used to produce cross-
sectional slices.
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64. Number of Detectors
 In a study comparing the dose from multislice CT scanners, Moore et al. (2006) demonstrated
that the measured radiation dose is inversely proportional to the number of detector rows.
 They showed that there is a trend that as the number of detector rows increases from 4 to 8 to
16 detector rows, the dose decreases with both standard and near identical technique
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65. STRATEGIES FOR REDUCING DOSE
1. Use CT Only When Clinically Indicated
2. Beam Filter
3. Adjusting mAs
4. Automatic Tube Current Modulation
5. Avoid Increasing kVp
6. Increased Pitch
7. Limit the Use of Thin Slices
8. Limit Repeat Scans
9. Newer Reconstruction Methods 9/21/2021 65

66. Dose Reduction Strategies
 General Strategies are based on the assumption that the CT scanner radiation dose level and image
quality fall within the manufacturer specification and other general quality criteria
 Proper patient positioning on a CT scanner table is the first step to assure use of dose saving
options.
 Higher noise images can occur when patients are not well centered in the scan field of view (FOV).
By positioning the body at isocenter, the need to increase mA to compensate for the noise is
eliminated.
66

67. Adjusting mAs
 adjust mAs to suit individual patient size
 Small bodies require a lesser dose, and large bodies require a greater dose.
 Use ATCM
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68. Avoid Increasing kVp
 Increasing the x-ray tube potential increases both the radiation dose and penetration of the x-rays
through the body.
 In general, increases beyond 120 kVp should be avoided, except when imaging obese patients.
 However, an increase in kVp could be accompanied by a reduction in tube current to offset the
increased dose.
 There is superior enhancement of iodine at lower tube potential which improves the differentiation
of vascular pathology.
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69. Limit the Use of Thin Slices
 Using a large number of thin adjacent CT slices results in 30% to 50% more radiation dose to
the patient than using fewer thicker slices to scan the same anatomy.
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70. Optimization of CT protocols
 CT technique should be tailored to the patient and his/her body habitus.
 CT technique should be monitored and controlled to ensure that the dose is as low as reasonably
achievable (ALARA principle):
 Reduction of peak KvP: relative organ dose reductions.
 Tube current adjustment (mAs): AEC software
 Adjust pitch: high pitch, decrease in dose
 Reduce the number of “multi-phasic” scans with contrast.
 Limit the scan range. 70

71. Automatic Tube current modulation
 ATCM is a generic name for any
technique aimed at optimizing dose
utilization by adjusting the tube
current in real-time to
accommodate differences in
attenuation due to patient
anatomy, shape, and size. The tube
current may be modulated as a
function of projection angle (Fig.
1a), longitudinal location along the
patients (Fig. 1b), or both (Fig. 1c).
 Hence data acquired through body
part having less attenuation can be
acquired with substantially less
radiation without negatively
affecting the final image noise .
71

72. 9/21/2021 72
(AEC) system: a) patient size, b) z-axis (AEC), c)
rotational/diagonal (AEC), d) combined (AEC)
Source.

73. 9/21/2021 73
Positioning the patient at isocenter is essential.
Otherwise, the attenuation calculated based upon
an off-centered Topogram will not be accurate. If a
patient is centered too low and the tube is
positioned under the table (Fig. 5a), the patient
appears wider on the Topogram than he/she actually
is (Fig. 5b), thus misinforming the CARE Dose4D
algorithm regarding the patient width. Similarly,
centering too high will misinform CARE Dose4D as to
the patient size (Fig. 5c).

74. Automatic Exposure Control (AEC)
 Adapt the tube current to patient attenuation and reduce the dose about 20- 40% under the proper
specification of image quality
74
MANUFACTURER AEC TRADE NAME
GE Auto mA/ smart mA
TOSHIBA Sure exposure
SIEMENS Care dose 4D
PHILIP Dose Modulation (DOM)

75. Cardiac CT
 ECG based tube current modulation is important dose reduction tool.
75

76. Advanced Iterative Reconstruction.
 Iterative recon. accurately modulates the system geometry incorporating physical effect such as beam
spectrum, noise & scatter.
 It Handles incomplete data set better than FBP.
 Hence, no. of projection/view can be significantly reduced without compromising image quality.
 There are a number of commercially-available algorithms (in 2019):
 iterative reconstruction in image space (IRIS), Sinogram-affirmed iterative reconstruction (SAFIRE)-
Siemens
 iDose4- Philips
 adaptive iterative dose reduction (AIDR) - Toshiba
 Adaptive statistical iterative reconstruction (AsIR) , Veo (model-based iterative reconstruction) – GE
76

77. 77

78. 78

79. 79
A–C, Low-dose CT scan obtained at 120 kVp, 3.75-mm
slice thickness, and CT dose index (CTDI) of 8 without
adaptive statistical iterative reconstruction (A) has more
image noise in liver than low-dose CT scan with adaptive
statistical iterative reconstruction (B) and routine-dose
CT scan (140 kVp; 3-mm slice thickness; CTDI, 22) (C). B
and C have nearly identical image quality.

80. 9/21/2021 80

81. Photon Counting Detector
 In contrast to the conventional EIDs, which integrate the energy levels of all detected
photons, PCDs count the number of individual photons that exceed a specified energy level.
 Signal from each interaction is individually processed
 Improves the SNR due to elimination of electronic noise.
 Benefit of SNR Improvement can be directly used to reduce patient dose.
81

82. Photon Counting Detector
82
Diagrams show detector types. A, In conventional energy-
integrating detector, an incident x-ray photon is converted into a
shower of visible light photons in a scintillator. Visible light hits an
underlying light sensor, where it generates positive and negative
electrical charges. B, In photon-counting detector, the x-ray
photon is absorbed in a semiconductor material, where it
generates positive and negative charges. Under the influence of a
strong electric field, the positive and negative charges are pulled
in opposite directions, generating an electrical signal.

83. 9/21/2021 83
Graph illustrates dose efficiency for low count
rates. A photon-counting detector has a
constant dose efficiency as a function of count
rate. For an energy-integrating detector, the
electronic noise becomes significant at very low
count rates, reducing dose efficiency.

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87. 2007
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91. Radiation Awareness
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92. Ten Steps to Wise and Gentle CT Protocols
1. Establish the appropriateness or justification guidelines for CT scanning.
2. Create clinical indication–specific CT protocols for different anatomic areas of the body. Such protocols
help tailor CT dose and image quality on the basis of specific clinical indications.
3. Ensure that the patient is positioned appropriately in the gantry iso-center.
4. Use AEC techniques to adapt tube current.
5. Use voltage adjustment on the basis of clinical indication and patient size.
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93. Ten Steps to Wise and Gentle CT Protocols
6. Select appropriate detector and section collimation on the basis of the clinical indication.
7. Avoid over-scanning beyond the anatomic area of interest.
8. Limit the use of multiphasic CT protocols. Multiphasic CT should be performed only when it is expected
to help in lesion detection or characterization.
9. Use IR techniques to improve image quality and allow dose reduction.
10. Assign a specific personnel team to be responsible for protocol review and dose monitoring.
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94. Summary
Radiation
Unit
Exposure
Absorbed Dose
Equivalent Dose
Effective Dose
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CT
Dose
Descriptors
MSAD
CTDI100
CTDIW
CTDIVOL
DLP
SSDE
STRATEGIES
FOR
REDUCING
DOSE
Clinical Indication
(Justification)
Exposure
Technique(ALARA,
Optimization)
Slice width &
Spacing, pitch
Filtration
Collimation
ATCM
Detector Efficiency
Noise Reducing
Recon. Algorithm

95. Conclusion
 Although the absolute risk of performing CT is not definitively known, the team of
radiologists, physicists, technologists, and CT technology manufacturers must optimize
available technology to reduce the potential risks from ionizing radiation while maintaining
the ability to perform high-quality imaging that allows clinical questions to be answered
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97.  Bushberg., 2012. Essential Physics of Medical Imaging, The. Lippincott, Williams & Wilkins, 3rd ed
 Romans, L. Computed tomography for technologists, 2nd ed
 Seeram, E. Computed tomography: Physical Principles, Clinical Applications, and Quality, 4th ed
 Bushong, S. Radiologic science for technologists 11th ed
 Goldman, Lee W. “Principles of CT: Radiation Dose and Image Quality.” Journal of Nuclear Medicine
Technology, Society of Nuclear Medicine, 1 Dec. 2007, tech.snmjournals.org/content/35/4/213.
 Parakh A, Kortesniemi M, Schindera ST. CT radiation dose management: A comprehensive optimization
process for improving patient safety. Radiology. 2016;280(3):663–73.
 Goo HW. CT radiation dose optimization and estimation: An update for radiologists. Korean Journal of
Radiology. 2012;13(1):1.
 Kalra MK, Sodickson AD, Mayo-Smith WW. CT radiation: Key concepts for gentle and wise use.
RadioGraphics. 2015;35(6):1706–21.
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Reference

98. Reference
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100. Question?
 What are the dose descriptors used in CT?
 What are the factors effecting CT dose?
 How can you optimize radiation dose in CT?
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