Online Appendix for the following October 20 JACC article by 1tjy8J

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									Online Appendix for the following May 2 JACC article



TITLE: Radiation Exposure of Computed Tomography and Direct Intracoronary
Angiography: Risk Has its Reward




AUTHORS: Pat Zanzonico, PhD, Departments of Medical Physics and Radiology

(Nuclear Medicine Section), Memorial Sloan Kettering Cancer Center, New York,

New York, Lawrence N. Rothenberg, PhD, Departments of Medical Physics and

Radiology (Nuclear Medicine Section), Memorial Sloan Kettering Cancer Center,

New York, New York, H. William Strauss, MD, Departments of Medical Physics and

Radiology (Nuclear Medicine Section), Memorial Sloan Kettering Cancer Center,

New York, New York




APPENDIX



Technical Aspects of Radiation Exposure

Except in the very early days of diagnostic radiology, when radiation precautions
virtually were nonexistent, radiation safety concerns have focused on stochastic (or
statistical) (1,2) effects. The rapidly increasing use of lengthy interventional
procedures has led to renewed concern regarding high-dose deterministic effects (3–
8). At the relatively low energies (typically <150 keV) and limited penetrabilities of
diagnostic X-rays, radiation doses are highest at the beam entrance point, and the
most likely deterministic effects therefore are skin damage, manifesting as erythema
or, in severe cases, ulceration. The threshold absorbed doses for deterministic skin
damage associated with acute irradiation, i.e., 200 cGy for transient erythema to 2,000
cGy for dermal ulceration (7,8), are at least two orders of magnitude greater than
those associated with either CT or noninterventional CCA. Accordingly, as addressed
by Coles et al. (9) in this issue of the Journal, the relevant radiogenic risks associated
with MSCT and CCA are stochastic risks, namely, carcinogenesis (10–15).

Direct measurement of the fluoroscopy skin entrance dose with dosimeters, although
the most accurate, has been superseded by an indirect method, the dose-area product
(DAP) (5,10,14,15). The standard DAP meter uses an air ionization chamber mounted
on the face of the X-ray tube collimator. The ionization chamber measures exposure
(i.e., the total electric charge produced by X- and gamma-rays per unit mass of air),
which is then converted to absorbed dose (i.e., the energy deposited per unit mass of
matter) using the “f factor,” or exposure-to-dose conversion factor (f = 34.9 to 36.8
Gy/C/kg or 0.90 to0.95 rad/R for water or soft tissue) (16). The dose-area product
reflects the absorbed dose over the entire X-ray field and is a function of the field size
and the exposure at the collimator; it usually is expressed in units of Gray-centimeter
squared (Gy-cm2). This value is sometimes recast as the “air-kerma-area product.”
“Kerma” is defined as the ratio

dEtr
dm

where dEtr is the sum of the initial kinetic energies of all the charged ionizing particles
liberated by uncharged ionizing particles (including photons) in matter and dm is the
mass of matter in which the charged ionizing particles were liberated (17–19). For air,
this quantity generally is referred to as “air kerma” or “free air kerma.” The measured
DAP (or, alternatively, the air kerma-area product) is independent of distance from
the focal spot because the X-ray field area varies directly and the exposure rate varies
inversely as the square of the distance from the X-ray focal spot to the measurement
point. A given DAP reading can result from a high dose over a small field or a low
dose over a large field. The stochastic risk can be assumed to be approximately
equivalent under these two conditions, and DAP measurements have thus been used
to estimate total stochastic risk (6). For selective contrast coronary angiography, DAP
varies from 6 to 109 Gy-cm2, with most reported values lying in the 30- to 60-Gy-cm2
range (20). In a novel study, Katritis et al. (21) used thermoluminescent devices
(TLDs) mounted on a specially designed catheter advanced to the right or left sinus of
Valsalva to directly measure the absorbed dose to the coronary arteries in CCA. A
linear relationship between the DAP and the coronary artery dose, DAP = 1,273
cm2·dose (mGy), was thus derived; the constant of proportionality is presumably
dependent on the system and the parameter. In another noteworthy study, Kuon et al.
(22) reported a 66% reduction in DAP (from 37.1 to 12.9 Gy-cm2) by limiting of
cinegraphic runs, systematic use of low-level fluoroscopy, and blind positioning of
the region of interest and resulting avoidance of oblique positions.

For MSCT, the basic radiation dose parameter is the computed tomography dose
index (CTDI), which defined as the integral under the exposure or absorbed dose
profile along the patient’s longitudinal (z) axis for a single tomographic image
(10,14,15). This parameter is scanner-specific, and its value generally is provided by
the manufacturer. The maximum of the radiation dose profile is termed the “peak
dose.” The volume CTDI (CTDIvol) is derived from the CTDI and is the average dose
delivered to a scan volume (vol) for a specific examination (10,14,15). The 100-mm
CTDI (CTDI100) is the integral under the exposure or absorbed dose profile along a
100-mm length of the patient’s longitudinal (z) axis (10,14,15). The weighted 100-
mm CTDI (CTDIw) is the weighted average of the CTDI100 measurements at the
center and periphery of a dose-measurement phantom (10,14,15):

CTDIw  [2/3 CTDI100 (p) + 1/3 CTDI100 (c)]·f (1)

where CTDI100 (p) = the CTDI100 at the periphery (p) of a cylindrical phantom,
CTDI100 (c) = the CTDI100 at the center (c) of a cylindrical phantom, and f = the
exposure-to-absorbed dose conversion factor = 34.9 to 36.8 Gy/C/kg, or 0.90 to 0.95
rad/R, for water or soft tissue.

The CTDIw thus reflects the average absorbed dose over the transverse (x and y)
dimensions of such a phantom and is an approximation of the average radiation dose
to the cross section of a patient. Measurements of the CTDI100 (p) and CTDI100 (c)
typically are performed using ionization chambers or TLDs positioned in a
commercially available soft tissue-equivalent polymethylmethacrylate (i.e., Plexiglas)
phantom cylindrical in shape and either 16 or 32 cm in diameter. Ionization chambers
actually measure exposure, which is then converted to absorbed dose using the
aforementioned f factor. On the other hand, TLDs, yield absorbed the dose directly.

Several parameters have been devised to estimate the dose associated specifically
with MSCT examinations (10,14,15). The multiple-scan average dose (MSAD) is the
average dose of the central scan of a MSCT examination. The MSAD is directly
related to the spatial separation of successive scans and, therefore, the “pitch,” or the
advance (or feed) of the patient table during a spiral CT examination. Pitch is now
rigorously defined as the distance (mm) of patient table advance in the longitudinal
(z) direction per gantry rotation divided by the total nominal scan length, and is thus a
dimensionless quantity (10,14,15). For MSCT systems, the total nominal scan length
includes all simultaneously acquired scans and is the distance in the z direction
spanned by all detector rows that are active during a scan. If the table advance during
one gantry rotation is less than the total nominal scan width (i.e., pitch <1), scans
overlap. Scan overlap, and therefore patient dose, increases as pitch decreases. The
dose-length product is defined as the product CTDIvol·total scan length; the total scan
length incorporates the number of scans and the length of each scan. Like the MSAD,
the dose-length product (mGy-cm) is thus a measure of the integral radiation dose of
an entire CT examination.

The ED is currently the most widely used and most rigorous measure of stochastic
risk. ED provides a single-value estimate of the overall stochastic risk (i.e., the total
risk of cancer and genetic defects) of a given irradiation whether received by the
whole body, part of the body, or a single or multiple individual organs (17–19):

ED  wT HT (2)
        T


=
       wT wR DT,R (3)
        R
    T

where wT is the weighting factor for tissue or organ T, a dimensionless quantity
representing the fraction contributed by tissue or organ T to the total stochastic risk
(i.e., the combined total risks of cancer or of demonstrable germ-cell mutagenesis)
due to a uniform, total-body irradiation and wR is the weighting factor for radiation R,
a dimensionless quantity selected to account for the differences in biological
effectiveness of different types of radiation and ranging from 1 for sparsely ionizing
X- and gamma-rays (including diagnostic X-rays) to 20 for densely ionizing alpha-
rays.

The ED is similar in concept to the effective dose equivalent, HE (introduced by the
International Commission on Radiological Protection and the NCRP) (1,19),
representing a single-value estimate of the net “harm” from any “low-dose” (e.g.,
diagnostic) exposure. However, the HE is related to the absorbed doses at individual
points within organs and thus is problematic to rigorously implement in practice. The
ED, in contrast, is based on average organ absorbed doses. The units of ED and HE
are the same: the signal intensity unit is the sievert (Sv) and the conventional unit is
the rem; 1 Sv = 100 rem and 1 rem equals 1 cSv (or 10 mSv). Neither the ED nor the
HE, however, are applicable to “high-dose” (e.g., interventional) exposures (23).

For estimation of EDs, organ-absorbed doses generally are either measured using
dosimeters (e.g., ionization chambers or TLDs) positioned in tissue-equivalent
anthropomorphic phantoms or calculated by Monte Carlo simulations in mathematical
anthropomorphic models (5,10,12–15,20). In the report by Coles et al. (9), a standard
adult hermaphrodite model, with a weight of 71 kg and height of 174 cm, was
mathematically modeled and used in conjunction with Monte Carlo analysis. The
average organ absorbed doses thus determined, the corresponding tissue weighting
factors (wT) and the radiation weighting factor(s) (wR = 1 for diagnostic X-rays) are
then substituted into Equation (3) to yield the ED. Of course, to the extent that the
size, shape, and composition of individual patients deviate from those of the phantom
(e.g., patients in the study by Coles et al. [9] typically were heavier and shorter than
the foregoing phantom), patients’ EDs will differ from the phantom-derived ED.

For MSCT coronary angiography, the scanner is set to record thin slices at high
resolution, requiring a high flux of X-rays. Coles et al. (9) found a mean effective
dose from MSCT coronary angiography of 14.7 mSv and that from selective coronary
angiography of 5.6 mSv. The dose these investigators measured is high compared to
publications by other investigators (Table 1), typically of the order of 10 mSv versus
1 mSV for CCA (5,12–15,20,21). For calcium-scoring MSCT, reported EDs range
from 2 to 4 mSv (24).
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Table 1. Reported Radiation Doses From Coronary CT Angiography


       Author            kVp           mA         Gated Acquisition            Pitch       ED in mSv (slices)

    Hoffmanni       120              240         Prospective ECG            0.2            8.6 (16 slice)

    Morinii         120              300         Prospective ECG            0.375          9.3, 11.3 (4 slice)

    Lauiii          140              250         Prospective ECG                           1.3-1.5 (4 slice)

    Hackeriv        120              500         Prospective ECG                           4.3 (16 slice)

    Hunoldv         120              400,                                   0.375          6.7-13 (4 slice)
                                     300,
                                     200



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