Добавил:
Sekretar
kiopkiopkiop18@yandex.ru
t.me/Prokururor I Вовсе не секретарь, но почту проверяю
Опубликованный материал нарушает ваши авторские права? Сообщите нам.
Вуз:
Предмет:
Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_3752_Библиотеки_им_академика_М_И_Перельмана
.pdf
26
Special Design Considerations
for Cardiac CT
Spatial Resolution
Spatial resolution is the product of multiple variables
including detector characteristics and collimation, sampling rate, and reconstruction methods. State of the art
scanners use thin 0.5–0.625 mm collimation. This reduces
volume averaging thus increasing image detail of small
objects, and allows reconstruction of isotropic datasets,
thereby preserving z-axis resolution and permitting multiplanar reformatting necessary for analysis of cardiac
anatomy.
GE Healthcare has focused their system development
on spatial resolution, developing a new detector scintillation material with a fast decay time (30 ns), minimal
afterglow (one-quarter that of conventional Gd
2
O 2 S crystals), and increased efficiency. This new detector is one
element of a multi-pronged approach to increasing image
resolution, which also includes new electronics that permit increased sampling, and a new iterative reconstruction algorithm [ 1 ] (Fig. 2.1 ). These innovations increase
spatial resolution by 3–4.5 line pairs per cm (lp/cm),
which is substantial given that current resolution of systems is approximately 12–15 lp/cm. The enhanced resolution is particularly useful for imaging stented and/or
calcified coronary arteries, while lowering radiation
requirements.
Temporal Resolution
In order to reconstruct an axial image, attenuation data from
multiple view angles that encompass half a rotation around
the patient are required. The amount of time needed to obtain
this data is the nominal, heart-rate independent temporal
resolution of the scanner. Hence, a gantry that requires
330 ms to make one complete rotation has a baseline temporal resolution of 165 ms. This is analogous to shutter speed:
just as shutter speed must be suffi ciently fast to capture
motion-free images of a moving subject, temporal resolution
must be suffi ciently fast to capture the data needed for image
reconstruction within the relatively brief period of cardiac
diastasis. If temporal resolution is insuffi cient, motion artifact occurs (Fig. 2.2 ). A temporal resolution of 165 ms is
relatively slow for cardiac imaging, so systems overcome
this by various means: faster gantry rotation, multi-cycle
reconstruction, and a second x-ray tube.
Multi-cycle reconstruction combines data from multiple
contiguous heart beats to complete the half-scan of views
(Fig. 2.3 ). This technique is heart-rate dependent: it can only
be used within certain ranges of heart rates, since the cardiac
cycle and gantry rotation must be asynchronous.
An approach adopted by Siemens in 2006 uses two x-ray
tubes mounted on the gantry at 90° angles; therefore, only a
quarter turn of the gantry is required to collect the 180° of
attenuation data. Hence, a most recent dual-source system
with a gantry rotation time of 250 ms has a baseline central
temporal resolution of approximately 66 ms. This obviates
Fig. 2.1 Increased sampling improves image quality. Axial images
of line-pair phantom demonstrate improved spatial resolution using
increased sampling density (on the right ) compared to standard sam-
pling frequency (on the left ). Note the improved edge detail and
resolution of the line pairs of the image derived from the high density sampling dataset (Reproduced with permission from Chandra
[ 2 ] )
R. Nakanishi et al.
https://t.me/medicina_free
27
the need for multi-cycle reconstruction, making the system
less susceptible to variation in heart rate, and providing high
heart rate independent temporal resolution. Several reports
demonstrated very good image quality [ 4 , 5 ] and, in a com-
parison of 64 MDCT and DSCT scans, the dual-source system produced better image quality at higher heart rates [ 6 ].
Image quality is still improved, however, by reducing heart
rate before scanning. The improvement of temporal resolution is not the only way to solve the issue for motion artifact.
As a post image processor, GE currently provides the motioncorrection algorithm (Snapshot Freeze; GE Healthcare)
which improves image quality using coronary motion artifact and supplements interpretability for diagnostic problems
of CAD [ 7 ].
Detector Coverage and Scan Speed
The several novel multi-slice CTs have achieved wider
z-axis coverage. The scan length is 16 cm, which is mostly
enough to scan the whole heart in a single heart beat,
avoiding breathing or misalignment artifact. In addition,
by increasing coverage or faster temporal resolution, scan
times have been recently reduced to just < 1 s, resulting in
the minimization of the radiation exposure. The Toshiba’s
Aquilion One scanner (Aquilion One Dynamic Volume
CT; Toshiba Medical System, Tochigi-ken, Japan) has 320
rows of detectors, with 0.5 mm collimation, 16 cm of coverage, and can perform an axial 1-beat acquisition of the
entire heart when the rate is well controlled. The GE revolution also demonstrated a wider coverage of 16 cm and
0.23 mm collimation, with gantry rotation 280 mm. The
Philips iQon (Philips Healthcare, Cleveland, OH), is a 256row scanner, which allows axial acquisition of the entire
heart in 2 beats.
Radiation Exposure
Traditional cardiac CT uses a helical scan mode and very low
pitch to perform retrospective ECG-gated reconstructions.
This delivers an undesirably high radiation exposure, so various methods have been designed to reduce radiation
exposure.
Fig. 2.2 Better temporal resolution improves image quality of moving
structures. 3D volume reconstructions of the right coronary artery in 3
CT scans representing 3 generations of CT scanner (4-, 16-, and
64-slice, from left to right , respectively), possessing temporal
resolution of 400, 250, and 180 ms from left to right , respectively.
Note that the vessel appears distorted and disjointed, due to motion
artifact, when temporal resolution is insuffi cient ( far left ), while
motion artifact is minimally present, and small anatomic details are
clearly imaged when temporal resolution is improved ( far right )
(Reproduced with the kind permission of Springer Science + Business
Media from Hurlock et al. [ 3 ] )
2 Cardiovascular Computed Tomography: Current and Future Scanning System Design
https://t.me/medicina_free
28
One of the fi rst was ECG-based tube current modulation,
which fl uctuates tube current in sync with the heart rate,
maintaining high tube current during diastasis and providing
the best image quality in that phase, then lowering tube current during other phases when high-resolution detail is not
required (Fig. 2.4 ). This results in a dose savings of
approximately 40 %.
A great advance in dose reduction was made with the
advent of prospective, ECG-triggered, axial cardiac CT
(Fig. 2.5 ). By keeping the x-ray tube off during most of the
scan, and only turning it on during diastasis, as triggered by
the ECG, radiation exposure is dramatically reduced. GE
fi rst published results using this method [ 9 ] with good clini-
cal results and radiation doses of 1–2 mSv [ 10 ], and this
approach is now offered by all manufacturers. Using a
64-slice system, the entire heart can be covered in 3–4
acquisitions; however this leaves the system vulnerable to
fl uctuations in heart rate and rhythm during the scan. Larger
coverage mitigates this vulnerability by reducing the number
of beats required for data acquisition. However, larger coverage lends to higher radiation doses, as the detectors are less
effi cient when having to expose a wider detector array.
Hence, Philips’ 128-row iCT system can cover the entire
heart in 2 beats, with a reported radiation exposure of
3–5 mSv. Toshiba’s 320-row Aquilion One scanner can
cover the entire heart in one beat, with a reported radiation
exposure of approximately 6 mSv. However, lower radiation
doses and dependable coronary scanning using 1-beat
acquisition requires heart rate reduction. At higher heart
rates (>65 bpm), optimal image quality requires a 2- or
3-beat acquisition (and use of multi-cycle recon) which
increases radiation exposure to approximately 13 (2-beat) to
19 (3-beat) mSv [ 11 , 12 ]. Acquiring with a wide exposure
window (“padding”) can ensure the capture of motion-free
data, but increases radiation exposure (Fig. 2.6 ). A narrower
exposure window can be used with preservation of image
quality if heart rate is controlled and reduces radiation
exposure.
High Pitch Helical Scanning
Radiation exposure is inversely proportional to pitch in
ECG-gated helical CT. Thus, increasing the pitch could
dramatically lower radiation requirements in cardiac
CT. In 2009, Siemens introduced an innovative scanning
method using a high pitch helical mode which takes advantage of the dual-source design [ 13 , 14 ] (Fig. 2.7 ). The high
pitch (3.2–3.4) would normally produce gaps in the attenuation data using a single-source system, but, in a dualsource system, these gaps are compensated for by gathering
data from the second detector. Scan time is less than one
second, with radiation exposures now reported less than
1 mSv. Initial studies performed on the fi rst-generation
dual-source system proved the feasibility of the method,
though the authors noted that these fi rst-generation systems are not suitable for this high pitch technique. High
pitch scanning with their new scanner, with faster gantry
rotation (250 ms) and larger coverage (96 rows), has been
reported to produce very good coronary image quality with
radiation exposure of <1 mSv [ 4 , 15 , 16 ], or even lower,
with <0.1 mSv [ 17 ]. Of note, in this initial investigational
phase, low heart rates (<60 bpm) are absolutely required,
as the entire data collection takes place within a 250–
270 ms acquisition window of one heart beat; hence
z
z
t
t
180°
45°
45°
45°
45°
Fig. 2.3 Multi-cycle reconstruction. Single cycle recon ( a ): The dura-
tion of the acquisition window ( gray bar ) is approximately equivalent
to one-half the gantry rotation time, since this is the time required to
obtain 180° of attenuation data. Multi-cycle recon ( b ): When multiple
detector rows are present, the axial slice position in the z-position can
be imaged at multiple different times, using multiple, shorter, acquisition windows distributed across multiple contiguous beats. This data
can then be combined to provide 180° of attenuation data and used to
reconstruct the axial image. Temporal resolution of the image is
improved because the duration of the each acquisition window is
shorter (Reproduced with permission of Wolters Kluwer from Vembar
et al. [ 8 ] )
R. Nakanishi et al.
https://t.me/medicina_free
29
diastasis must be at least this long to provide motion free
data acquisition.
The GE Revolution can now utilize the new detector
and reconstruction system, along with a fast pitch algorithm, to lower radiation exposure by preserving image
quality while scanning with less radiation (<1 mSv),
because images are of higher resolution and lower noise
than would otherwise be achievable with a low-exposure
scan.
Future Directions
Future scanner designs are closely guarded industry secrets,
but some concepts have been openly discussed. Flat-panel
volume CT systems replace detector rows with a large area
detector, and provide coverage of the entire heart in one
axial acquisition and extremely high spatial resolution of up
to 26 lp/cm, comparable to invasive angiography [ 19 ].
However, the contrast resolution is inferior to that of multidetector CT, a high radiation exposure is needed to achieve
a suffi cient contrast-to-noise ratio, scintillation times are
slow and temporal resolution is insuffi cient for cardiac
imaging.
Dual energy scanning is being developed by all vendors
by various means, including rapid fl uctuation of tube energy,
stacked detectors, or the dual-tube design in which each tube
100 %
20 %
Time
Tube output
Fig. 2.4 ECG-based tube current
modulation reduces radiation
exposure. Once the desired
acquisition phase is determined,
based on heart rate, in this case
75 % phase, scanning proceeds
with maintenance of 100 % tube
output during that phase, while
tube current is reduced outside of
that phase. This can reduce
effective radiation dose by 40 %
or more; however, note that
images reconstructed from the
low tube output phases ( left
image ) will be excessively noisy,
and are usually considered
unusable for coronary
interpretation (Reproduced with
permission of Wolters Kluwer
from Vembar et al. [
8 ] )
Table move
75 % 75 %
Fig. 2.5 Prospective ECG-triggered acquisition method. Cardiac
rhythm is monitored while table remains stationary. When cardiac cycle
reaches pre-determined acquisition phase (in this case 75 %, diastolic,
phase), x-ray source is briefl y turned on (<1 s, blue bars ) and acquires
attenuation data of a length equivalent to the craniocaudal coverage of
the detector array, and is then turned off. The table is advanced almost
the length of craniocaudal coverage (minus a small amount of overlap),
and the process repeats again. Depending on the craniocaudal coverage
of the scanner, the entire heart can be scanned in 1–4 acquisitions
(Reproduced with the kind permission of Springer Science + Business
Media from Weigold et al. [ 18 ] )
2 Cardiovascular Computed Tomography: Current and Future Scanning System Design
https://t.me/medicina_free
30
emits photons of different energy levels. Dual energy CT
could take the fi eld to a new level of diagnostic power if it
can be used for refi ned tissue characterization, such as differentiating plaque characteristics. This is diffi cult to do in
the coronary arteries, and initial studies have not yielded any
breakthroughs, but at the least it has the potential to enhance
coronary lumen visualization in heavily calcifi ed vessels by
differentiating calcium and iodine [ 20 , 21 ] (Fig. 2.8 ). Dual-
energy CT may also yield new applications for non-coronary
imaging, such as imaging of myocardial perfusion [ 22 ].
Low heart rate
75 % 75 %
High heart rate
Table move
Table move
40 % 40 %
a
b
Fig. 2.6 Prospective CCTA phase selection and padding. Acquisition
phase and padding depend on heart rate: For low heart rates ( a ), target
late diastole (75 % phase) for prospective acquisition, but if heart rate is
high ( b ) (>75 bpm), target end-systole (40 % phase) which is more
likely to produce motion-free images. By restricting acquisition to the
minimum exposure window ( dark gray ), radiation dose is minimized,
but ability to reconstruct adjacent cardiac phases is obliterated.
Widening the acquisition window (“padding,” light gray shading)
allows reconstruction of a small number of additional phases, which
can help interpretation of motion artifact, but increases the radiation
dose (Reproduced with the kind permission of Springer
Science + Business Media from Weigold et al. [ 18 ] )
83 ms
Fig. 2.7 Method of high-pitch coronary CCTA. Single source helical
CT requires a pitch ≤ 1.5; a faster pitch results in gaps in the data ( left
panel ). A dual-source system can scan at a higher pitch (up to 3.2),
using the second x-ray source-detector system to fi ll in what would otherwise be gaps in the data. Since the table speed is so high, the entire
heart can be imaged in a fraction of a second, from data derived from a
single heartbeat ( right panel ). If the image acquisition is triggered
appropriately to synchronize acquisition with diastasis, and the heart
rate is suffi ciently low, motion-free images can be obtained with a very
low radiation dose (Reproduced with permission of Elsevier from
Achenbach et al. [ 14 ] )
R. Nakanishi et al.
https://t.me/medicina_free
31
Conclusions
Since the turn of the twenty-fi rst century, there has been
an explosion in technological development of cardiac CT
systems, with concomitant gain in reliability and accuracy, especially of coronary imaging. The theme has been
one of progressively improved spatial and temporal resolution, reduced scan time, and, most recently, a focus on
driving down radiation exposure. In appropriately selected
patients, using careful technique, it is easily achievable to
perform a cardiac CT using less radiation exposure than
that of a standard chest CT or an invasive coronary angiogram. The goal of future systems will be to make this
more widely achievable in a larger group of patients without requiring patient selection or stringent patient prep.
Given the history of rapid technological advancement, we
can expect to see this goal achieved in the near future.
References
1. Thibault JB, Sauer KD, Bouman CA, Hsieh J. A three-dimensional
statistical approach to improved image quality for multislice helical
CT. Med Phys. 2007;34:4526–44.
2. Reproduced with permission from Chandra N. CT sampling tech-
nology [white paper]. Waukesha: GE Healthcare; 2008.
3. Hurlock GS, Higashino H, Mochizuki T. History of cardiac com-
puted tomography: single to 320-detector row multislice computed
tomography [Review]. Int J Cardiovasc Imaging. 2009;25:31–42.
4. Hell MM, Bittner D, Schuhbaeck A, et al. Prospectively ECG-
triggered high-pitch coronary angiography with third-generation
dual-source CT at 70 kVp tube voltage: feasibility, image quality,
radiation dose, and effect of iterative reconstruction. J Cardiovasc
Comput Tomogr. 2014;8:418–25.
5. Achenbach S, Ropers D, Kuettner A, et al. Contrast-enhanced
coronary artery visualization by dual-source computed tomography – initial experience. Eur J Radiol. 2006;57:331–5.
aa
cd
Fig. 2.8 Dual energy CT. Dual energy CT can be used to characterize
tissue: Vessel cross sectional images derived from high- (140 keV) and
low-energy (90 keV) attenuation profi les ( a and b , respectively) demon-
strate the infl uence of photon energy on photoattenuation. The contrastfi lled lumen ( arrow ) exerts greater photoattenuation on low- energy
photons, and hence appears brighter in ( b ), while calcifi cation ( aster-
isk ) appears high-density in both images. In the subtracted image ( c ),
dense calcifi cation has been “removed.” By adding in the low- energy
(90 keV) attenuation data to this subtracted image, depiction of the
lumen edge is enhanced (because of reduced contrast blooming), and
visualization of a small side branch ( arrowhead ) is improved ( d )
(Reproduced with permission of Wolters Kluwer from Boll et al. [ 23 ] )
2 Cardiovascular Computed Tomography: Current and Future Scanning System Design
https://t.me/medicina_free
32
6. Achenbach S, Ropers U, Kuettner A, et al. Randomized comparison of 64-slice single- and dual-source computed tomography
coronary angiography for the detection of coronary artery disease.
JACC Cardiovasc Imaging. 2008;1:177–86.
7. Leipsic J, Labounty TM, Hague CJ, et al. Effect of a novel vendorspecifi c motion-correction algorithm on image quality and diagnostic accuracy in persons undergoing coronary CT angiography
without rate-control medications. J Cardiovasc Comput Tomogr.
2012;6:164–71.
8. Vembar M, Walker MJ, Johnson PC. Cardiac imaging using
multislice computed tomography scanners: technical considerations. Coron Artery Dis. 2006;17:115–23.
9. Hsieh J, Londt J, Vass M, Li J, Tang X, Okerlund D. Step-and-shoot
data acquisition and reconstruction for cardiac x-ray computed
tomography. Med Phys. 2006;33:4236–48.
10. Earls JP, Berman EL, Urban BA, et al. Prospectively gated transverse coronary CT angiography versus retrospectively gated helical
technique: improved image quality and reduced radiation dose.
Radiology. 2008;246:742–53.
11. Hoe J, Toh KH. First experience with 320-row multidetector CT
coronary angiography scanning with prospective electrocardiogram
gating to reduce radiation dose. J Cardiovasc Comput Tomogr.
2009;3:257–61.
12. Steigner ML, Otero HJ, Cai T, et al. Narrowing the phase window
width in prospectively ECG-gated single heart beat 320-detector row
coronary CT angiography. Int J Cardiovasc Imaging. 2009;25:85–90.
13. Hausleiter J, Bischoff B, Hein F, et al. Feasibility of dual-source
cardiac CT angiography with high-pitch scan protocols.
J Cardiovasc Comput Tomogr. 2009;3:236–42.
14. Achenbach S, Marwan M, Schepis T, et al. High-pitch spiral acquisition: a new scan mode for coronary CT angiography. J Cardiovasc
Comput Tomogr. 2009;3:117–21.
15. Gordic S, Desbiolles L, Sedlmair M, et al. Optimizing radiation dose
by using advanced modelled iterative reconstruction in high- pitch
coronary CT angiography. Eur Radiol. 2016;26(2):459–68.
16. Layritz C, Schmid J, Achenbach S, et al. Accuracy of prospectively ECG-triggered very low-dose coronary dual-source
CT angiography using iterative reconstruction for the detection of coronary artery stenosis: comparison with invasive
catheterization. Eur Heart J Cardiovasc Imaging. 2014;15:
1238–45.
17. Schuhbaeck A, Achenbach S, Layritz C, et al. Image quality of
ultra-low radiation exposure coronary CT angiography with an
effective dose <0.1 mSv using high-pitch spiral acquisition and raw
data-based iterative reconstruction. Eur Radiol. 2013;23:
597–606.
18. Weigold W, Olszewski M, Walker MJ. Low-dose prospectively
gated 256-slice coronary computed tomographic angiography. Int
J Cardiovasc Imaging. 2009;25 Suppl 2:217–30.
19. Gupta R, Cheung AC, Bartling SH, et al. Flat-panel volume CT:
fundamental principles, technology, and applications.
Radiographics. 2008;28:2009–22.
20. Andreini D, Pontone G, Mushtaq S, et al. Diagnostic accuracy of
rapid kilovolt peak-switching dual-energy CT coronary angiography in patients with a high calcium score. JACC Cardiovasc
Imaging. 2015;8:746–8.
21. Barreto M, Schoenhagen P, Nair A, et al. Potential of dual-energy
computed tomography to characterize atherosclerotic plaque:
ex vivo assessment of human coronary arteries in comparison to
histology. J Cardiovasc Comput Tomogr. 2008;2:234–42.
22. Ruzsics B, Schwarz F, Schoepf UJ, et al. Comparison of dualenergy computed tomography of the heart with single photon emission computed tomography for assessment of coronary artery
stenosis and of the myocardial blood supply. Am J Cardiol.
2009;104:318–26.
23. Boll DT, Hoffmann MH, Huber N, Bossert AS, Aschoff AJ, Fleiter
TR. Spectral coronary multidetector computed tomography angiography: dual benefi t by facilitating plaque characterization and
enhancing lumen depiction. J Comput Assist Tomogr. 2006;30(5):
804–11.
R. Nakanishi et al.
https://t.me/medicina_free
© Springer International Publishing 2016
M.J. Budoff, J.S. Shinbane (eds.), Cardiac CT Imaging: Diagnosis of Cardiovascular Disease,
DOI 10.1007/978-3-319-28219-0_3
Radiation Dosimetry and CT Dose
https://t.me/medicina_free
Reduction Techniques
Kai H. Lee
Abstract
Increasing radiation exposure to the population from widespread use of multi-row detector
CT necessitates efforts to limit the x-rays applied in CT procedures. Options are available
on CT scanners to modulate the patient dose. This chapter describes the metrics of radiation
dose, and the influence of various technical parameters in the scan and dose modulation
options on patient dose and image quality. With that information, practitioners of CT will
be better prepared to optimize the scan protocols to reduce the patient dose without sacrificing the image quality necessary for interpretation.
Keywords
CT radiation metrics • Dose modulation • Scanning parameters
3
Introduction
The ability of modern multi-detector CT scanners with
sub- millimeter resolution, sub-second rotation time, and large
volume imaging has resulted in widespread use of
CT. However, the widespread use of CT has also raised concerns about the risks of radiation to patients. The National
Council on Radiation Protection, NCRP Report No. 160, in
2006 [1] reported that radiation exposure to the United States
population due to medical sources increased more than seven
times in the 20 years between 1986 and 2006. According to
the NCRP report, CT constituted about 10 % of the diagnostic
examinations that utilize x-rays in 2006, it contributed to
nearly 50 % of the population dose. The use of CT continued
to rise since publication of NCRP Report 160. Reports in sci-
K.H. Lee, PhD
Associate Professor of Clinical Radiology,
Department of Radiology, Keck School of Medicine,
University of Southern California,
1200 North State Street, Los Angeles, CA 90033, USA
e-mail: kailee@usc.edu
entific and lay journals found fewer than three million CT
examinations done in 1980, which rose to 62 million in 2007,
and to 80 million at the end of 2014 [2–4]. Based on our
current knowledge of radiation biology, the deleterious effects
of radiation are cumulative and medical radiation is increasingly a significant contributor to the amount of radiation accumulated in a person’s lifetime [5, 6]. The risk of cancer from
radiation exposure is especially worrisome to children and
young women who received multiple CT examinations early
in their life. For example, studies found that one CT examination of the female chest gives as much radiation as 10 mammograms to each breast [7]. In addition to radiation from CT,
cardiac patients may be exposed to radiation from nuclear
medicine perfusion studies and coronary angiography.
Therefore, the practitioners of CT must be constantly aware
of the risks of radiation, and strive toward applying the lowest
dose to the patient consistent with the clinical study.
One of the difficulties confronting the clinicians when
evaluating the radiation safety of a CT procedure is the plethora of terms used to quantify the amount of radiation given to
the patient. Thus, this chapter sets out on two aims. The first
aim is to explain the fundamental concepts of radiation
dosimetry associated with CT scans. The second aim is to
describe the scanning techniques and available technologies
to reduce radiation dose to patients.
33
34
11
yG
-
W.
()=()
()
=
()
×
11
K.H. Lee
https://t.me/medicina_free
Fundamentals of Radiation Dosimetry
Absorbed Dose
The damaging health effects of radiation are commonly
called radiation effects. Our current knowledge of radiation
effects was derived from animal studies, atom bomb
survivors, and victims of radiation accidents [6]. There were
many contributing factors to the observed radiation effects,
but the outcomes ultimately depended on the amount of
radiation received. The radiation dose, or dose in short, is
the amount of energy deposited by the ionizing radiation per
unit mass of tissue. Radiation dosimetry is the field of study
which measures and quantifies the radiation dose. Of the
bewildering number of terms to quantify radiation dose,
there are three related terms fundamental to radiation
protection. The three terms are the absorbed dose, the
equivalent dose, and the effective dose.
The International System (SI units) of radiation dose
measurements is the Gray [8]. One Gray (Gy) of radiation is
defined as 1 J of energy deposited in 1 kg of tissue, i.e.,
joule
Dose Gy
()
1
=
kg tissue.
1
The Gray is a large dose of radiation. When evaluating
radiation dose from CT examinations, a sub-unit of the Gray
is used. The Gray sub-unit commonly used to quantify
radiation dose from CT examinations is the milli-Gray
(mGy). One mGy is one thousandth of a Gray. That is,
3
mG
0
y.=
Equivalent Dose
The severity of biological damage depends not only on the
dose of radiation absorbed in tissue, but also on the type of
radiation absorbed. For example, 1 mGy of neutron radiation produces far greater tissue damages than 1 mGy of
x-rays. We therefore need a unit of measurement that
accounts for both the quantity of radiation absorbed in the
tissues and the effectiveness of the absorbed radiation in
producing biological damages. The equivalent dose was
devised to incorporate both the physical factors and the
effectiveness of the absorbed radiation in producing biological damage. The equivalent dose is a biological scale for
measuring the deleterious health effects of radiation. The
same biological damage is expected for the same equivalent
dose regardless of the type of radiation absorbed in tissue.
The equivalent dose (H) is measured in Sievert (Sv).
A sub-unit of Sievert is the milli-Sievert (mSv). One
Table 3.1 Biological damage weighting factors
Type of radiation Weighting factor
X- and gamma rays, electrons, positrons 1
Protons 2
Neutrons 10
Alpha particles 20
milli-Sievert equals to one thousandth of a Sievert. The
equivalent dose in mSv is calculated by multiplying the
radiation absorbed dose D in mGy by a quality factor Q, also
called the biological damage weighting factor Wr [9], for the
type of radiation absorbed in tissue, i.e.,
HmSv DmGy
´
r
The biological damage weighting factors for four types of
radiation are given in Table 3.1.
The values of Wr are proportional to the density of ionization created by the incident radiation along its path of travel
in tissue. For x-rays, gamma rays, beta particles, and electrons from radioactive materials, the density of ionization
created in tissue is relatively low. The weighting factor Wr
equals 1. Thus, when working with x-rays from CT, the
equivalent dose and absorbed dose are numerically equal,
i.e., 1 mSv =1 mGy. For neutrons, the weighting factor
Wr= 10. The equivalent dose for 1 mGy of neutron absorbed
dose equals
HmSv mGy
mSv.
=
10
0
The above example shows that neutrons are ten times more
damaging to the human body than x-rays for the same
absorbed dose.
In summary, the absorbed dose in mGy is the quantity of
radiation energy deposited per unit mass of tissue. The
equivalent dose in mSv is a measure of biological damage
equal to the absorbed dose modified by a weighting factor
according the relative effectiveness of the absorbed radiation
to produce biological damage.
Given this metric to quantify radiation, Table 3.2 lists
the average equivalent dose received annually to the
total body by workers in various occupations [10–13].
The table also gives the natural background radiation
and the regulatory limits on radiation exposure for reference. It is interesting to note that the transcontinental
flight crews who are not classified as occupational radiation workers receive an annual equivalent dose from
the cosmic rays higher than the nuclear medicine technologists who routinely handle radioactive materials on
the job.
EH
=∑
ii
3 Radiation Dosimetry and CT Dose Reduction Techniques
https://t.me/medicina_free
35
Table 3.2 Typical annual whole-body radiation dose
mSv
Nuclear medicine technologists 1.7
Airline flight crews between Los Angeles and
New York City
Nuclear power plant workers 2.3
Intervention radiologists 18
Cardiologists (catheterization) 16
Natural background radiation at sea level 2.5
Regulatory limit on the occupational workers 50
Regulatory limit on the general public 1
2.2
Radiation Effects to the Patient
The goal of protecting patients undergoing radiological procedures is to prevent the occurrence of deterministic effects,
and to minimize the risk of stochastic effects. Deterministic
effects are radiation induced somatic injuries. Examples of
radiation induced somatic injuries are skin erythema, epilation, and cataracts. The FDA has documented severe deterministic effects such as hair loss and skin necrosis on patients
from CT and prolonged fluoroscopy guided procedures.
Deterministic effects are preventable. There is ample clinical
evidence of threshold doses below which such effects are not
seen. Once the threshold dose is reached, severity of the
injury increases with the radiation dose.
Stochastic risks are probabilistic occurrences of carcinogenesis and genetic mutations. The current concept of stochastic risks assumes no threshold dose for its occurrence.
Exposure to any amount of ionizing radiation is harmful. The
probability of the occurrence of stochastic effects, rather
than its severity, increases in direct proportion to the radiation absorbed dose. This is in contrast to deterministic effects
in which the severity increases as the dose increases above
the threshold. This concept of stochastic risks is called linear
no threshold model, and is the basis of radiation protection
regulations.
Effective Dose
Our knowledge of stochastic risks is based on data collected
from total body exposure to radiation. If we wish to estimate
the stochastic risks to a person after a partial body exposure
such as a chest CT, we must translate the chest CT dose to an
equivalent whole-body dose in order to utilize the database
for risk estimates. The effective dose was devised for this
purpose. The effective dose is the dose of radiation that when
given uniformly to the whole-body will produce the same
stochastic risk as the dose of radiation delivered to only a part
of the body, such as the chest dose from a cardiac CT study.
Table 3.3
Tissue/organ Weighting factor, Wr
Gonads 0.20
Bone marrow, colon, lung, stomach 0.12
Bladder, breast, liver, esophagus, thyroid 0.05
Skin, bone surface 0.01
Tissue/organ sensitivity
The effective dose translates a partial body exposure to an
equivalent uniform dose of radiation delivered to the total
body. The purpose for calculating the effective dose is to provide a common denominator for the assessment of stochastic
risks using our database of whole-body exposures. Unlike
the absorbed dose and the equivalent dose, the effective dose
is not a physically measurable quantity. The effective dose is
an imaginary total body dose. It is calculated from the
absorbed dose given to any region or regions of the body.
If we wish to estimate the stochastic risks from CT of the
chest, we must translate the partial body irradiation to an
equivalent whole-body dose in order to utilize the database
for risk estimates. To do so, a mathematical model is used to
compute the doses to other organs resulting from radiation
scattered from CT of the chest. These computed organ doses
are then multiplied by a risk factor according to the suscepti-
bility of each organ to radiation. Summation of the product
of these computed organ doses and their associated risk
weighting factor is called the effective dose. That is, the
effective dose E is computed using the equation
w
Where E is the effective dose
Hi is the dose equivalent to a given organ
wi is the risk weighting factor for that organ.
The effective dose is thus a weighted sum of the computed equivalent doses to all organs in the body. Table 3.3 is
a partial list of weighting factors for different body organs
published in the International Commission on Radiation
Protection Report 103 [14].
One may interpret the effective dose as a calculated
equivalent dose of radiation given to the entire body that
would be required to produce the same risk of cancer and
genetic damage as a dose of radiation delivered to a localized
region of the body, such as in a CT examination. In other
words, the risk from a part of the body exposed to a given
dose of radiation is the same as the total body uniformly
receiving the effective dose. The effective dose is an extrapolated whole-body dose from a partial body dose. As such, the
effective dose is a computed value rather than a physically
measurable quantity. The effective dose is calculated to serve
as a common denominator for comparing stochastic risks
Соседние файлы в папке Библиотека им академика М.И. Перельмана