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Radiation Dose From CT of the Heart

CYNTHIA H. MCCOLLOUGH, PhD

INTRODUCTION

The issue of radiation dose from X-ray computed tomogra- phy (CT) has received much attention recently in both the popu- lar media and scientific literature (1–5). This is in part due to the fact that the dose levels from CT typically exceed those from conventional radiography and fluoroscopy, and that the use of CT continues to grow. Thus, CT contributes a significant por- tion of the total collective dose from ionizing radiation deliv- ered to the public from medical procedures. It is important, therefore, that physicians ordering or performing these exami- nations have an understanding of the dose delivered from a cardiac CT, as well as how that amount of radiation compares to those from other imaging procedures that use ionizing radia- tion.

HOW TO DESCRIBE THE DOSE FROM A CT EXAMINATION: CTDI AND DLP

CT dose descriptors, the basic tools required for understand- ing radiation dose in CT, have been in existence for many years, yet continue to be refined as multidetector-row CT (MDCT) evolves. The primary measured value is known as the CT Dose Index (CTDI) and represents the integrated dose, along the z axis, from one axial CT scan (one rotation of the X-ray tube) (6–8) (Fig. 1). Typically, a 100-mm long ionization chamber is used for routine measurements. Thus, the subscript 100 is used to denote the measurement length. All other CT dose descriptors are derived from this primary measured value. It is important to note that the CTDI is always measured in the axial scan mode, and that doses for helical scan modes are calculated from the axial information. The equipment used to measure CTDI is shown in Fig. 2.

The CTDI varies across the field of view. For example, for body CT imaging, the CTDI is typically a factor or two higher at the surface than at the center of the field of view. The average CTDI across the field of view is given by the weighted CTDI (CTDIw), where CTDIw = 2/3 CTDI(edge) + 1/3 CTDI(center) (9–10). Figure 3 gives the typical relative distribution of dose in the head and body phantoms. CTDIw is a useful indicator of scanner radiation output for a specific kVp and mAs. CTDIw is reported in terms of absorbed dose to air (9–10).

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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 radiation dose pro- files from consecutive rotations of the X-ray source. This is accomplished with use of a dose descriptor know as the Vol- ume CTDIw (CTDIvol), where

CTDIvol = [(N × T)/ I ] × CTDIw and

N = the number of simultaneous axial scans per X-ray source rotation

T = the thickness of one axial scan (mm) I = the table increment per axial scan (mm) (10).

In helical CT, the ratio of the table travel per rotation (I) to the total nominal beam width (N × T) is referred to as pitch (10–11). Hence

CTDIvol = (1/pitch) × CTDIw.

So, whereas CTDIw represents the average radiation dose over the x and y directions, CTDIvol represents the average radiation dose over the x, y, and z directions. This provides a single CT dose parameter, based on a directly and easily mea- sured quantity, which represents the average dose within the scan volume for a standardized (CTDI) phantom (10). CTDIvol is a useful indicator of the dose for a specific exam protocol, because it takes into account protocol-specific information such as pitch. Its value may be displayed prospectively on the con- sole of newer CT scanners, although it may be mislabeled on some systems as CTDIw. Recent consensus agreement on these definitions is reflected in newer scanner software releases (10).

Thus, CTDIvol estimates the average radiation dose within the irradiated volume of a CT acquisition. The SI units are milliGray (mGy). It does not indicate, however, the total energy deposited into the scan volume. Its value remains unchanged whether there are 20 or 40 scans acquired.

To better represent the overall energy (or dose) delivered by a given scan protocol, the dose can be integrated along the scan length to compute the dose-length product (DLP), where

DLP (mGy-cm) = CTDIvol (mGy) × scan length (cm) (9).

The DLP reflects the total energy absorbed (and thus the potential biological effect) attributable to the complete scan acquisition. Thus, a limited abdomen CT might have the same

From: Contemporary Cardiology: CT of the Heart:

Principles and Applications

Edited by: U. Joseph Schoepf © Humana Press, Inc., Totowa, NJ

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Fig. 1. Computed Tomography Dose Index (CTDI) is the integral under the radiation dose profile from a single axial scan.

Fig. 2. Typical equipment used for measuring CT dose index.

Fig. 3. Typical dose distributions (%) across the image field-of-view.

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Table 1

Sample Volume CT Dose Index (CTDIvol)

and Dose Length Product (DLP) Values for Common CT Exams

Chest Abdomen Abdomen and pelvis

Peak kilovotage (kVp) 120 120 120

Tube current (mA) 200 300 300

Exposure time (s) 0.5 0.5 0.5

Detector configuration (N × T) 4 × 5 mm 4 × 5 mm 4 × 5 mm

Table index per rotation (I) 15 mm 15 mm 15 mm

Pitch (I/N × T) 0.75 0.75 0.75

Reconstructed scan width (mm) 5 5 5

Scan length (cm) 40 20 40

CTDIvol (mGy) 12.0 19.1 19.1

DLP (mGy × cm) 480 382 764

CTDIvol as an abdomen and pelvis CT, but the latter exam would have a greater DLP, proportional to the greater z extent of the scan volume.

Table 1 demonstrates the differences in CTDIvol and DLP for typical body CT exams. The values are for demonstration only; they can vary by scanner model, vendor, and image qual- ity requirements. Note that a change in technique (mAs/rota- tion) affects the CTDIvol, while a change in acquisition length (at the same technique) is reflected by the DLP.

In cardiac CT, the anatomic scan length is relatively con- stant (typically 12 cm); thus, the variability in CTDIvol and DLP is primarily a result of differences in scanner output and scan acquisition parameters. Tables 2 and 3 provide the scan acqui- sition parameters, CTDIvol, and DLP for coronary calcification imaging and coronary angiography. Data are provided for an electron beam CT (EBCT) system as well as for MDCT sys- tems from two different manufacturers.

AUTOMATIC EXPOSURE CONTROL

It is technologically feasible for CT systems to adjust the X-ray tube current (mA) in response to variations in X-ray intensity at the detector (12–13), much as fluoroscopic X-ray systems adjust exposure automatically. This capability, in vari- ous implementations, is now available commercially on MDCT systems in response to wide interest from the radiology com- munity. Some systems adapt the tube current based on changes in attenuation along the z axis, others adapt to changes in attenuation as the X-ray tube travels around the patient. The ideal is to combine both approaches with an algorithm that

“chooses” the correct tube current to achieve a predetermined level of image noise.

With regard to cardiac CT, the radiation dose for a retrospec- tively gated exam, where the X-ray tube is kept continuously on throughout the acquisition, can be dramatically reduced if the tube current is reduced during portions of the cardiac cycle Table 2

Scan Acquisition Parameters, Volume CT Dose Index (CTDIvol),

and Dose Length Product (DLP) for Coronary Calcification Imaging (12-cm scan length)

EBCT MDCT1 MDCT1 MDCT2 MDCT2

Data acquisition method Prospective Prospective Retrospective Prospective Retrospective triggering triggering gating triggering gating

Peak kilovotage (kVp) 130 120 120 120 120

Tube current (mA)a 630 140 100 150 150

Exposure time (s) 0.1 0.36 0.5 0.33 0.5

Detector configuration (N × T) 1 × 3 mm 4 × 2.5 mm 4 × 2.5 mm 4 × 2.5 mm 4 × 2.5 mm

Table index per rotation (I) 3 mm 10 mm 3.75 mm 10 mm 3.75 mm

Pitch (I/N × T) 1 1 0.375 1 0.375

Reconstructed scan width (mm) 3 2.5 3 2.5 2.5

CTDIvol (mGy) 3.5 4.6 12.5 4.7 20.3

DLP (mGy × cm) 42 55 150 56 243

Effective doseb (mSv) 0.7 0.9 2.6 1.0 4.1

a mA can be increased in multidetector-row CT (MDCT) scanners for larger patients to avoid an increase in image noise. The mA values for MDCT 1 and MDCT 2 were provided by the respective manufacturers, and do not necessarily produce an identical level of image noise.

b Effective Dose estimate, with k = 0.017 mSv × (mGy × cm)–1. This value is averaged between male and female models (see text and ref. 5).

EBCT, electron beam CT.

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Table 3

Scan Acquisition Parameters, Volume CT Dose Index (CTDIvol)

and Dose Length Product (DLP) for Coronary Angiography (12-cm Scan Length)

EBCT MDCT1 MDCT2

Data acquisition method Prospective Retrospective Retrospective

triggering gating gating

Peak kilovotage (kVp) 130 120 120

Tube current (mA)a 630 300 300

Exposure time (s) 0.1 0.5 0.5

Detector configuration (N × T) 1 × 3 mm 4 × 1 mm 4 × 1.25 mm

Table index per rotation (I) 2 mm 1.5 mm 1.9 mm

Pitch (I/N × T) 0.66 0.375 0.375

Reconstructed scan width (mm) 3 1.25 1.25

CTDIvol (mGy) 5.3 46 55

DLP (mGy × cm) 64 547 662

Effective doseb (mSv) 1.1 9.3 11.3

a mA can be increased in multidetector-row CT (MDCT) scanners for larger patients to avoid an increase in image noise. The mA values for MDCT 1 and MDCT 2 were provided by the respective manufacturers, and do not necessarily produce an identical level of image noise.

b Effective Dose estimate, with k = 0.017 mSv × (mGy × cm)–1. This value is averaged between male and female models (see text and ref. 5).

EBCT, electron beam CT.

that are not likely to be of interest for the reconstructed data.

Thus, in addition to modulation of the tube current based on patient attenuation, the tube current can be modulated by the ECG signal. Since cardiac motion is least during diastole and greatest during systole, the projection data are least likely to be corrupted by motion artifact for diastolic-phase reconstruc- tions. Accordingly, the tube current is reduced during systole.

Dose reductions of approx 50% have been reported using such a strategy (14). The implementation of these and other dose- reduction strategies is expected industry-wide over the next several years, in response to the strong concern about the radia- tion dose from CT.

EFFECTIVE DOSE

It is important to recognize that the potential biological effects from ionizing radiation depend not only on the radiation dose, but also on the biological sensitivity of the tissue or organ system irradiated. A 100-mGy dose to an extremity would not have the same potential biological effect (detriment) as a 100-mGy dose to the pelvis (15). Effective dose (E) is a dose descriptor that reflects this difference in biologic sensitivity (16–17). It is a single dose parameter that reflects the risk of a nonuniform exposure in terms of an equivalent whole-body exposure. The units of E are milliSieverts (mSv).

Although the concept of effective dose has some limitations when applied to medical populations, it does facilitate the com- parison of biological effect between diagnostic exams of dif- ferent types (16,17). Published values of E per DLP (9) allow convenient estimates of E based on the DLP value provided at the CT scanner console. The use of E facilitates communication with patients regarding the potential harm of a medical exam that uses ionizing radiation. For example, when a patient inquires, “What dose will I receive from this exam,” an answer in the units of mGy or mGy × cm will not likely answer the more

fundamental, but perhaps unspoken, question, “What is the likelihood that I will be harmed from this exam.” Characteriz- ing the radiation dose in terms of E and comparing that value to some meaningful level—for instance, one year’s E from naturally occurring background radiation—better conveys to the patient the relative potential for harm from the medical exam. Table 4 provides typical values of E for several common imaging exams, as well as the annual level of background radiation in the US (approx 3.6 mSv).

It is important to remember, however, that E describes the relative whole-body dose for a particular exam and scanner, but is not the dose for any one individual, as E calculations use many assumptions including a mathematical model of a “stan- dard” human body that does not accurately reflect any one individual. Effective dose is best used to optimize exams and compare risks between proposed exams. It is a broad measure of risk, and as such should not be quoted with more than one or two significant digits.

Specific values of E can be calculated using several different software packages (17), which are based on the use of data from one of two sources: the National Radiological Protection Board (NRPB) in the UK (18) or the Institute of Radiation Protection (GSF) in Germany (19). To minimize controversy over differ- ences in E values that are purely the result of calculation meth- odology and data sources, a generic estimation method was proposed by the European Working Group for Guidelines on Quality Criteria in CT (9), where E is estimated from the non- controversial value of DLP: E = k × DLP, where the values of k are dependent only on the region of the body being scanned (head, neck, thorax, abdomen, or pelvis) (Table 5). The values of E predicted by DLP and the values of E estimated using more rigorous calculation methods are remarkably consistent, with a maximum deviation from the mean of approx 10–15%. Hence, the use of DLP to estimate E appears to be a reasonably robust

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method for estimating E. The effective doses for several car- diac CT examinations, based upon values of DLP, are given in Tables 2 and 3.

SUMMARY

The fundamental dose parameter in CT, the CTDI, is mea- sured at edge and center locations in an acrylic phantom for a given kVp, mAs, and scan width, and reported for a given exam protocol as the volume CTDI (CTDIvol) in units of mGy. This value represents the average dose in a standard acrylic phan- tom for a given exam protocol. Another relevant dose param- eter is that of effective dose, which is given in units of mSv. The effective dose is a single dose parameter that best represents the radiation detriment corresponding to a given exam protocol.

Neither of these parameters is an estimate of radiation dose to any one individual, but rather should be used to optimize and compare exam protocols. E can be estimated with good accu- racy from the DLP, which is equal to the CTDIvol multiplied by the total scan length (in cm).

Techniques to modulate the tube current as a function of patient attenuation or the time within the cardiac cycle are important innovations that will reduce the dose from cardiac CT by at about a factor of two. Hence, coronary artery calcium examinations, which currently have E values between 1 and 4 mSv, may be able to be conducted using retrospective gating techniques and an E of less than 2 mSv. CT coronary angiog- raphy, which currently requires an E of approx 10 mSv with retrospectively gated MDCT may be performed with approx 5 mSv if ECG tube current modulation is applied. These effective dose values are of similar magnitude to those from a chest CT examination (5–7 mSv) or a conventional (diagnostic) coro- nary angiogram (3–5 mSv).

Finally, as cardiac CT image quality, diagnostic accuracy, and availability all continue to improve, there will be more and

more publications regarding the clinical efficacy of cardiac CT as compared to alternate imaging modalities. Discussions of patient safety and societal cost will include discussions of the radiation dose from the various procedures. To ensure that these discussions are accurate, it is essential that the dose informa- tion associated with specific techniques be reported in an accu- rate and complete fashion using standardized terminology.

REFERENCES

1. Brenner DJ, Elliston CD, Hall EJ, Berdon WE. Estimated risks of radiation-induced fatal cancer from pediatric CT. AJR 2001;176:

289–296.

2. Donnelly LF, Emery KH, Brody AS, et al. Minimizing radiation dose for pediatric body applications of single-detector helical CT:

strategies at a large children’s hospital. AJR 2001;176:303–306.

3. Haaga JR. Radiation dose management. AJR 2001;177:289–291.

4. Nickoloff EL, Alderson PO. Radiation exposure to patients from CT: reality, public perception, and policy. AJR 2001;177:285–287.

5. Pierce DA, Preston DL. Radiation-related cancer risks at low dose among atomic bomb survivors. Radiation Research 2000;154:178–186.

6. Shope TB, Gagne RM, Johnson GC. A method for describing the doses delivered by transmission x-ray computed tomography. Med Phys 1981;8:488–495.

7. American Association of Physicists in Medicine. Standardized methods for measuring diagnostic x-ray exposures. Report no. 31.

AAPM, New York, 1990.

8. Nagel HD. Radiation exposure in computed tomography. Frankfurt:

COCIR, 2000.

9. European guidelines for quality criteria for computed tomography.

Luxembourg: European Commission, 2000.

10. International Electrotechnical Commission. Medical Electrical Equipment. Part 2–44: Particular Requirements for the Safety of X-ray Equipment for Computed Tomography. IEC publication No.

60601-2-44 Amendment 1.

11. McCollough CH, Zink FE. Performance evaluation of a multi-slice CT system. Medical Physics 1999;26:2223–2230.

12. Gies M, Kalender WA, Wolf H, Suess C, Madsen M. Dose reduction in CT by anatomically adapted tube current modulation I: simulation studies. Medical Physics 1999;26:2235–2247.

13. Kalender WA, Wolf H, Suess C. Dose reduction in CT by anatomi- cally adapted tube current modulation II: phantom measurements.

Medical Physics 1999;26:2248–2253.

14. Jakobs TF, Becker CR, Ohnesorge B, et al. Eur Rad 2002;12:

1081–1086.

15. Committee on the Biological Effects of Ionizing Radiation. Health effects of exposure to low levels of ionizing radiation, BEIR V.

Washington, DC: National Academy, 1990.

16. International Commission on Radiological Protection (ICRP). 1990 Recommendations of the ICRP. Publication 60. ICRP, New York, NY, 1991.

Table 4

Effective Dose Values for Common Imaging Examinations

Examination Effective dose (mSv)

Head CT 1–2

Chest CT 5–7

Abdomen and pelvis CT 8–11

Selective coronary angiogram 3–5

Posterior-anterior and lateral chest X-ray 0.04–0.06 Average annual background radiation in the US 3.6

Table 5

Values of the Conversion Coefficient k for Use in Estimating Effective Dose (E) (mSv) From Dose Length Product (DLP)

(in mGy × cm) According to the Formula E = k × DLP (9) Anatomic region k (mSv × mGy–1× cm–1)

Head 0.0023

Neck 0.0054

Chest 0.017

Abdomen 0.015

Pelvis 0.019

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17. McCollough CH, Schueler BA. Calculation of effective dose. Medi- cal Physics 2000;27:828–837.

18. Jones DG, Shrimpton PC. Survey of CT practice in the UK. Part 3:

Normalised organ doses calculated using Monte Carlo techniques, NRPB-250. Oxon, United Kingdom: National Radiological Protec- tion Board, 1991.

19. Zankl M, Panzer W, Drexler G. The calculation of dose from exter- nal photon exposures using reference human phantoms and Monte Carlo methods. Part IV: Organ dose from computed tomographic examinations, GSF-Bericht 30/91. Neuherberg, Germany: GSF – Forschungszentrum fur Umwelt und Gesundtheit, Institut fur Strahlenschutz, 1991.

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