dose

The Physics of CT Dose

Richard Mather, PhD

Senior Manager, Clinical Science

Toshiba America Medical Systems, Inc.

For more than 30 years, Computed

Tomography (CT) has allowed physicians to

explore the human body in ways that were

previously impossible. The technology has

progressed from single-slice axial scans

that required hours to perform to modern

multidetector helical exams that can image

the entire body in a few seconds (Figure 1).

CT is now the modality of choice for a wider

range of clinical applications than all other

imaging modalities combined. However, all

the benefits of CT must be balanced against

the risks associated with x-ray exposure.

While the biological effects of high

doses of radiation are well known and

documented3,4,5,6,7, the effects of low doses,

such as those received from CT exams,

are not as clear8,9. Most of what we believe

about low-dose ionizing radiation comes

from information extrapolated from atomic

bomb survivors3,4, who received far greater

doses of radiation than a patient typically

Computed Tomography (CT) has been the fastest growing technology in medical imaging for more than

half a decade. Recent advances in Multi Detector CT (MDCT) have enhanced the number and range of CT

imaging applications available to physicians and their patients. These advancements have increased the

clinical utility of CT and over 62 million CT procedures are currently being performed annually in the United

States. With this increased use of CT has come a growing awareness of the potential risks from the radiation

dose associated with a CT scan1,2. In order to knowledgably balance clinical benefi ts gained from CT with

the potential risks of radiation exposure, it is critical to understand the factors that contribute to CT dose, as

well as how dose levels are estimated. Finally, it is essential to understand the ALARA principle (As Low As

Reasonably Achievable) of radiation use. ALARA requires the use of the lowest radiation dose that will yield

appropriate image quality for a particular patient to enable the correct clinical decision.

The Physics of CT Dose 1

Figure 1: 64-slice MDCT enables the rapid acquisition of thin-slice, volume datasets over large portions of the anatomy. 0.5 mm slices produce 350 micron isotropic resolution volumes, allowing images to be viewed in any orientation with no loss of spatial resolution.

2 The Physics of CT Dose

receives from a CT (Figure 2). From

these high dose effects, a conservative

model of radiation risk, known as the

linear no-threshold (LNT) model, was

developed10,11,12,13. The LNT model assumes

that the effects of high radiation doses can

be extrapolated to low doses. The model

also assumes that any dose, no matter how

small, has the potential to cause harm. At

low patient doses, the primary concern is

the possibility of developing cancer 10 to 20

years after the exposure14.

It is important to balance the potential

risk of a CT examination with the clinical

benefits: the dose used in a given exam

must be enough to deliver sufficient image

quality to answer the clinical question but as

low as possible to minimize the risk to the

patient. This approach is defined as the As

Low As Reasonably Achievable (ALARA)

principle of radiation use3,4,15.

Under the ALARA principle, it is important

to consider radiation dose in conjunction

with a defined level of acceptable image

quality. Failing to consider the image

quality at a particular exposure level can be

potentially dangerous: too low a dose may

compromise the necessary image quality

and lead to either misdiagnosis or even

the need to repeat the entire scan. If a

rescan is required, the dose will be higher

due to the need for a higher mA technique,

and the dose from the original scan will

be wasted. Thus, adhering to the ALARA

principle does not simply mean reducing

the dose for all exams. The “appropriate”

dose will vary with the clinical context.

In order to determine the appropriate

dose for a given patient or exam, the

clinician needs to define the clinical

question. For example, to facilitate liver

transplant surgery planning, CT is often

used to determine the size of the patient’s

liver. For this clinical task, only a very

low dose may be all that is required since

only large, well-defined structures need

to be captured and high image noise

will not significantly affect the ability to

determine an accurate three-dimensional

liver volume. On the other hand, a higher

dose would be appropriate if the clinical

task were to visualize a small, low-contrast

lesion on a large patient’s liver because

it could be easily hidden by noise16. The

key is to adjust the radiation dose to the

diagnostic task at hand, using the lowest

possible dose that will yield the needed

1.09

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Rela

tive

Risk

Effective Dose (mSv)

Relative Biological Risk

Typical CT (10 mSv)

Measured Risk

0 50 100 150 200 250

Figure 2: The biological risks from ionizing radiation are clear above 50 mSv of whole body dose based on the BEIR VII data. The risks below 50 mSv are not measured, but are extrapolated using the assumption of the Linear No Threshold model. A typical CT exam is on the order of 10 mSv.


information. For this reason, use of

ionizing radiation needs to be performed

within accepted appropriateness criteria

and under the guidance of an appropriately

trained physician and technologist.

A classic example of the ALARA principle

at work is in pediatric CT. Because of their

size and level of radiosensitivity, if children

were to be scanned with the same scan

techniques as adults they would actually

receive effective doses 6 to 7 times larger

than necessary to achieve the diagnostic

task17, 18. Software like Toshiba’s

SUREExposure automatically chooses the

correct dose based on the patient’s size

and ensures that the ALARA principle is

upheld for all patients: pediatric and adult.

DOSE EFFICIENCY AND LOW CONTRAST

DETECTABILITY

Dose efficiency is the ability of a CT scanner

to achieve a particular level of image quality

at the lowest possible dose (Figure 3). The

best way to evaluate dose efficiency is to

look at the system’s ability to discriminate

objects with small differences in Hounsfield

Units (HU) from the background. Low-

contrast detectability (LCD) is closely

correlated with image noise as random

variation in an image will make small

objects that are similar in HU value to the

background hard to see (Figure 4). Since

noise depends on radiation dose and LCD

depends directly on image noise, LCD is

an excellent measure of a system’s overall

dose efficiency. The most dose efficient CT

system will be able to resolve the smallest

low-contrast object at the lowest dose19,20,21.

FACTORS AFFECTING PATIENT DOSE

AND IMAGE QUALITY

There are a number of scan parameters and

patient attributes that influence the dose and

image quality in a CT exam22,23,24,25. Some are

user controlled (e.g. kV, mAs, pitch). Other

factors are inherent to the scanner (e.g.,

detector efficiency, geometry). Still others

are patient dependent (e.g., patient size,

anatomy scanned). All these parameters

are interrelated. A solid understanding of

how each parameter relates to the others

and affects both dose and image quality is

essential to maintaining ALARA.

USER-CONTROLLED FACTORS

kV (kiloVolts) – Tube potential, or kV, is the

amount of voltage between an x-ray tube’s

anode and cathode. It determines the

energy of the x-rays being emitted. Higher

Figure 3: Low contrast resolution of different CT manufacturers from Product Data Sheets (as of September 2007). Toshiba can achieve 2 mm low contrast resolution with about the same or with less dose than other manufacturers need to resolve 3, 4, and 5 mm low contrast objects. Similarly, the dose Toshiba needs to resolve those 3, 4, and 5 mm low contast objects is 40-60% less than the other manufacturers.

Figure 4: Low Contrast Detectability (LCD) increases as the noise decreases.

30

25

20

15

10

5

0

Dose

(mGy

)

Low Contrast ResolutionDose Comparision

a

b

c

Toshiba

Object Size (mm)

5 4 3 2


energy x-rays have a greater probability than

lower energy x-ray of passing through the

body and creating signal at the detector.

With all else being equal, higher kV means

less noise. However, when these high-

energy x-rays are absorbed by the body,

they deposit more energy than lower

energy x-rays and, therefore, contribute

more to patient dose. For the same scan

parameters, changing the kV from 120 to

135 increases the dose by about 33%.

The image noise is reduced since the dose

is higher and more photons are reaching

the detectors, but the tissue contrast is

compromised as well. The main reason for

using higher kVs, such as 135 kV, is to get

better penetration for very large patients or

dense anatomy such as bone.

mAs (milliAmpere-seconds) – The tube

current, or mA, determines the number of x-

rays the tube produces. Combined with the

gantry rotation time, this represents the total

x-ray output of the tube per rotation, or mAs

(pronounced “mass”). Changing the mAs is

the most common method of adjusting dose

and noise level. For example, cutting the

mAs in half will reduce the patient dose by

a factor of 2, but also halve the number of x-

rays reaching the detector. Since fewer x-rays

are being detected, the image noise will be

increased and the LCD will be diminished26.

CT Pitch and Helical Pitch – Beam pitch

is defined as the distance the table travels

in a rotation divided by the total active

detector width in the Z direction. Helical

pitch is the same except it is divided by the

individual channel thickness rather than

the nominal total collimation. By either

definition, the higher the pitch, the faster

the table moves through the x-ray beam

and, consequently, the lower the dose to

the patient. It is important to note that

multislice scanners are different from single-

slice scanners with respect to pitch. Due to

the amount of redundant data with 64-slice

scanners, good reconstruction algorithms

can prevent the slice sensitivity profile from

broadening with pitches between 1 and 2.

However, as the pitch increases, there is

less projection data for a given slice and the

noise will increase with higher pitch values.

Therefore, while increasing the pitch lowers

the patient dose, it may also increase noise,

meaning one must increase the mA as the

pitch increases to maintain image quality.

Higher pitch value may also result in more

helical artifact. Therefore, the main clinical

Figure 5: Illustration of the effect of pitch on patient dose. For the same mA and rotation time, higher beam pitch values (>1) spread the X-rays out and reduce dose while lower beam pitch values (<1) concentrate the X-rays, increasing the dose.

Width of Helical Path (From Collimation)

Center of Helical Path

Direction ofContinuousPatient Transport

O

Z, mmt, sec

Pitch=1




O

Z, mmt, sec

Pitch=2Extended Helical




O

Z, mmt, sec

Pitch=0.5Overlapping Helical

reason to scan with a high pitch value is to

reduce the total scan time.

mAseff

(Effective mAs) – The mAseff

is

simply the mAs divided by the pitch. It is a

common, if imprecise, practice to use mAs

as a surrogate for dose (DLP, as discussed

later, is a better value to use). However, since

pitch affects the patient dose, mAs by itself

does not completely represent the number

of x-rays entering the patient (Figure 5). By

dividing the standard mAs by the beam pitch,

a value that is proportional to the patient dose

on a given scanner is derived. While using

mAseff

as a dose surrogate works adequately

well for comparing relative patient doses on a

single scanner, it does not work for comparing

doses from two different scanners. Many

aspects of the scanner’s design play a role in

the total dose delivered and the same mAseff

on two different scanners does not equate

to the same patient dose. The best way

to compare dose between scanners is

to examine the low-contrast detectability

(Figure 6).

Collimation – With multislice scanning, there

are many combinations of slice width and

number of slices that may be used to acquire

the scan volume. With all collimations on

multislice systems, the actual x-ray beam is

slightly wider than the nominal beam width

(i.e., number of detector rows multiplied by

detector element width). This is to ensure

that the detectors on the edge of the array

receive uniform x-ray coverage, resulting in

a small amount of unused radiation called

penumbra. Typically, the total amount

of penumbra is the same regardless of

the nominal beam width. Therefore, with

larger beams, the extra radiation from the

penumbra is a smaller percentage of the

overall useful beam width. Therefore, for the

best image quality and dose efficiency, the

thinnest slices covering the entire detector is

ideal (Figure 7).

Acquired and Reconstructed Slice Width –

The acquired slice width is governed by

the multislice detector configuration and

determines the minimum image width that

can be reconstructed. For example, images

acquired at a detector configuration of 16

x 1 mm can be reconstructed at 1 mm

or greater while images acquired at 16 x

0.5 mm can be reconstructed as thinly as

0.5 mm slices. Since one of the factors

contributing to image noise is the number

of photons used to create the image, thicker

reconstructed images have less noise with all


Figure 6: Effect of noise on low-contrast detectability. Whether due to lower dose on the same scanner or lower dose effi ciency between scanners, the system noise has signifi cant effects on the scanner’s ability to differentiate low-contrast objects.

Figure 7: Aquilion detectors from 4- to 64-slice. All Aquilion scanners have a consistent detector design manufactured to image with as thin as 0.5 mm slices and cover up to 32 mm per gantry rotation. The 4-, 8-, and 16-slice systems have 0.5 mm central detector elements with 1mm elements on the outside edges. The 32- and 64-slice detector is completely comprised of 0.5 mm elements. This design allows fl exability of scanning and uniformity of image quality across all the systems.


other factors being equal. However, thicker

slices have less z-axis resolution. Figure 8

shows the effect of imaging with different slice

thicknesses. Even with overlapping slices, the

thicker images do not allow the image clarity

available with the 0.5 mm acquired slices. All

of Toshiba’s multislice scanners, from 4 to 64

slice, offer 0.5 mm slices, the thinnest in the

industry. The Aquilion’s Quantum detector

consists entirely of 0.5 mm channels. On

the Aquilion 64, most exams are acquired in

the 64 x 0.5 mm configuration for the best

image quality and the lowest total penumbra

contribution. To reduce the image noise, it is

always possible to reconstruct thicker slices

from the thin, 0.5 mm acquired dataset.

Therefore, scanning with the thinnest possible

slices avoids partial volume artifacts (when a

single voxel contains more than one tissue),

allows excellent multiplanar reformats (MPR)

and 3D renderings (Figures 8 and 9), and

retains the most flexibility in reconstruction

without resorting to a higher radiation dose.

In light of the dose effects of collimation, the

optimal image quality and dose efficiency are

achieved with the thinnest slices covering the

entire detector.

Reconstruction Kernel – As part of the

reconstruction process, the acquired data is

filtered by the reconstruction kernel. Since

the reconstruction kernel plays a large role in

determining spatial resolution, it has a great

effect on the amount of noise in the image

and consequently, on the dose needed

for a given level of image noise. There is

a large variety of reconstruction kernels to

chose from. Sharper kernels lead to better

in-plane spatial resolution (at the cost of

greater noise), while smooth kernels reduce

the image noise at the expense of some

in-plane resolution. The choice of kernel is

made based on the clinical need: most body

kernels are relatively soft to reduce the image

noise and enhance contrast resolution,

while lung kernels tend to be sharp to better

identify the tiny airways.

SCANNER-INHERENT FACTORS

Detector efficiency – No single hardware

aspect of a scanner has more influence on

dose than the efficiency of the detection

system. The detector’s ability to catch the x-

ray, convert it to light, transmit that light, and

convert it to an electrical signal with minimal

loss defines the overall efficiency of the

detector. More efficient detectors result in

lower patient dose for a given level of image

quality. In order to create the highly efficient

illumination detector material, Toshiba’s

Figure 8: Effect of axial slice width on image reformats. When thick axial images are acquired or reconstructed, the multiplanar reformat image quality that can be derived from them deteriorates. A) shows an MPR made from 5 mm images with no overlap. B) shows an MPR made from 5 mm images with 4 mm of overlap. C) shows 0.5 mm images with no overlap. Note the slight increase in clarity for (B) and the signifi cant increase in clarity for (C). Thinner axial slices result in better reformatted images.

Figure 9: Effect of axial slice thickness on polyp detection. The same dataset was reconstructed using 4, 3, 2, and 1 mm axial slices. With the 4 mm and 3 mm datasets, the polyp is not distinguishable from a diverticular fold. At 2 mm, it becomes visible, but poorly defi ned. At 1 mm, the defi nition is greatly enhanced.

A B C

scientists invented a new method of sintering

the Gadolinium Oxysulfide (GOS) ceramic27.

With this new method, only Praesodynium

is added to the base ceramic which keeps

the light output high and allows the material

to be accurately machined to 0.5 mm

slices with clean, sharp edges. Toshiba’s

Quantum Detector system, based on the

illumination material, is the only system

that has maintained 0.5 mm slices from the

4-slice scanner all the way up to the 64.

illumination’s low afterglow and fast decay

times allow fast scanning and rotation times

without image artifacts while its high light

output contributes to the system’s industry

leading low contrast detectability.

Filtration – X-ray tubes, as the name implies,

produce the x-rays that are detected to

create the CT image. These x-rays, upon

leaving the tube, have a spectrum of

energies that range from very low energy

up to a maximum energy equal to the kV

(Figure 10). As discussed earlier, lower

energy x-ray photons do not penetrate

through the body as well as higher energy

ones. In fact, the lowest energy photons

will not pass through the body at all. Since

image formation relies upon photons getting

through the body and being picked up by the

detectors, these lowest energy photons only

contribute to patient dose. Therefore, all CT

scanners add a certain amount of filtration

outside the x-ray tube to block the low energy

x-rays. There is, however, a tradeoff involved

when choosing the amount of filtration: in

the process of removing low energy x-rays,

some desirable, medium and high energy x-

rays will be removed as well, thus decreasing

the overall output of the tube. This means

that higher mA values are needed to realize

a given image noise level. Furthermore, it

is the medium energy x-rays that provide

the best soft tissue contrast. Thus, heavier

filtering will compromise the system’s LCD.

Therefore, a CT system needs enough

filtration to block the lowest energy photons

but not so much as to lose the ability to

optimally distinguish low-contrast anatomy.

Geometry – It is a common misconception

that a CT scanner’s geometry (i.e., focus to

isocenter and focus to detector distances)

and focal spot dimensions play a significant

role in patient dose. In reality, the focal spot

dimensions only affect the overall dose by a

few percent. It is the ratio of the scanner’s

focus to isocenter and focus to detector

distances combined with the physical size of

the detector that determines the geometry’s


Initial Spectrum in a Vaccuum

Characteristic X-Ray Peak

Max kV

Initial Spectrum in Air

Filtered SpectrumNum

ber o

f X-

rays

X-ray Energy

Figure 10: Typical bremsstrahlung spectrum with and without fi ltration. The initial spectrum (orange), in a vacuum, steadily decreases up to the kV setting. in air, the lowest energy x-rays are fi ltered out, but a large number of low energy x-rays remain (blue). Many of the lowest energy x-rays will not penetrate a patient and only contribute to dose. All CT scanners employ a certain amount of additional fi ltration to “harden” the beam and remove the low energy x-rays (green).


role. From the diagram in Figure 11, it is

clear that as long as the detection system

size increases in proportion to the distance

from the focal spot, there is no difference

in mAs technique (i.e., dose) needed to

achieve the same image quality between

the two geometries. There is actually a skin

dose advantage to a longer geometry since

the patient is further away from the x-ray

tube during scanning. Furthermore, longer

geometries are more resistant to scatter since

there is a higher probability that a scattered

photon will miss the detector entirely.

PATIENT FACTORS

Patient Size – The size of the patient plays a

large role in the total dose absorbed for the

same technique. Figure 12 shows the effect of

patient thickness on the absorbed dose.

With smaller patient sizes, such as in a pediatric

case, the dose can be two to three times higher

than on an adult when using the same imaging

technique. However, with smaller patient sizes,

less kV and mAs are needed to achieve the

same image quality. Therefore, it is essential

to tailor the kV and mAs to the patient size to

minimize the dose28,29,30.

Body Part Scanned – Different organs

in the body have different sensitivities to

radiation. For example, the stomach is

more sensitive to radiation than the liver,

which is more sensitive to radiation than

bone. This means that for the same

absorbed dose, there is more risk if the

abdomen is scanned versus anatomy that is

mostly bone, such as the lower leg. This is

largely due to the rate of cellular replication

in these organs5.

Age – The risk of developing cancer from

a CT exam decreases with age. The latent

period, or time between radiation exposure

and the clinical onset of cancer, is on the

order of 10 years or more. Furthermore,

younger patients organs are more

radiosensitive in general, due to the rapid

rate of cell division and growth. Therefore, it

is critical in young patients to keep the dose

as low as possible while maintaining the

necessary image quality.

DOSE MEASUREMENTS

In order to adequately compare doses

between different CT systems, it is

necessary to understand how dose is

measured and reported. While it would be

ideal to measure the actual radiation dose

to each patient undergoing a CT scan, such

measurements are not possible. Rather,

Focal Spot

Short Geometry

Long Geometry

ApertureAperture

Detection System Size

Figure 11: Schematic of the dose implications of different focus to detector geometries. X-rays spread out as they get further from the focal spot. However, if the detector size increases with the longer distance, the same number of x-rays are captured with either geometry.

3.50

3.00

2.50

2.00

1.50

1.00

0.50

0.00

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tive

CTDI

Phantom Diameter (mm)

50 100 150 200 250 300 350

Figure 12: Relationship of body size to dose for a given x-ray technique. Smaller patients have less mass attenuating the x-rays and therefore receive more dose for the same technique. For this reason, pediatric protocols should use much less mAs than adult protocols.

proxies for dose, or “dose indices” have

been established. These measurements

are straightforward and repeatable in

nature. However, it is important to note

that, with current technology, these values

are only proxies for dose, as they are based

on simplifying assumptions and phantom

measurements. The common uses and

misuses of these measurements are

discussed below.

CTDI – For CT, a dose proxy known as

Computed Tomography Dose Index (CTDI)31,

is estimated for the head and body using two

standard plastic phantoms. One phantom

is 16 cm in diameter to approximate the size

of a head, and the other phantom is 32 cm

in diameter to approximate the size of a torso

or abdomen. The main purpose of these

phantoms is to simulate the attenuation and

scatter from a patient in the scanner. The

phantoms each have several holes designed

to hold a 10 cm ion chamber for direct dose

measurements: one in the center of the

phantom and four near the edge at the 3,

6, 9, and 12 o’clock positions. By making

measurements in each of these holes with

a given protocol, it is possible to derive a

number, or “index”, that is representative

of the dose absorbed in the phantom. It

is important to note that since a patient is

neither perfectly cylindrical nor made of

plastic, the CTDI value does not represent

the actual dose to the patient. Rather, CTDI is

a useful tool in comparing the relative doses

between two protocols on a given scanner.

Furthermore, CTDI is typically measured

using a 10 cm ion chamber which has been

shown to miss the edges of the scatter tails,

especially for wider beam widths.

There are several versions of CTDI that are

discussed in the medical physics community,

but the two most common and most useful

are CTDIw and CTDIvol. CTDIw is the

weighted average of the measured doses in

the 5 phantom hole positions given by:

CTDIw represents the axial absorbed dose

to the phantom for a given protocol and is

only applicable to non-helical protocols.

For helical protocols, similar to the way that

mAseff

is used to account for the overlapping

acquisition, CTDIvol takes pitch into account:

It is important to understand that while CTDI

can be useful when comparing protocols on


0.5 mm high-resolution lung CT.Effective dose of 5.6 mSv.

3D volume rendered high-resolution lung CT.Effective dose of 5.6 mSv.


a given scanner, it is misleading to use it to

compare the dose delivered by two different

scanners. A common misuse of CTDI is to

equate the normalized version, nCTDI, with

scanner dose. nCTDI is the CTDI value

for a given mAs value, typically 100 mAs.

However, the measured CTDI for a given mAs

value depends on the system’s filtration: more

filtration means fewer x-rays coming out of the

tube. For a given mAs value, a heavily filtered

system will have a lower nCTDI value than a

system with less filtration. However, it will also

have significantly lower image quality as well

for that same mAs value. This will be evident

in the LCD measurement and the image

noise. In order to achieve the same signal-to-

noise ratio, the heavily filtered system will have

to increase its mAs value to match the image

quality of the system with less filtration. This

also means that scanners with high filtration

will require larger generators to achieve the

same level of image quality.

This dependence on beam filtration

illustrates the main reason that CTDI is not a

good metric for dose comparisons between

scanners. The main drawback to CTDI is

that there is no image quality component

to it and the only way to compare doses

between different scanner systems is by first

equalizing image quality. Since low-contrast

detectability combines aspects of both image

quality and dose into a single scanner metric,

it is a far better choice when comparing the

dose efficiency of two different scanners.

As stated above, nCTDI (per 100 mAs) is

frequently misused to compare the doses

delivered from different scanners. Since

nCTDI values are only for a given mAs

setting, they cannot be used to compare

actual patient doses. Different scanners use

different mAs values for the same clinical

task and image quality, due to differences

in filtration. Therefore, unless the image

quality is equalized between the systems, any

dose comparison between those systems is

meaningless. Scanner comparison reports,

such as those from ImPACT and ECRI, as

well as the manufacturer’s own product data

sheets, by convention report the scanner’s

nCDTI. While these values contain good

information about the tube output and beam

filtration, it is important to remember they

are not indications of the scanner’s dose

efficiency.

As an example, if scanner A has a CTDI per

100 mAs in the body of 10 mGy and scanner

B has a CTDI per 100 mAs of 8 mGy, we

Cardiac CTA.Effective dose of 8.6 mSv.

Abdominal aorta with runoff using SUREExposure.Effective abdominal dose of 4.7 mSv.

cannot yet say whether scanner A or B will

deliver less dose. If we determine the image

quality is equivalent when scanner A uses

150 mAs and scanner B uses 250 mAs,

then we could make a comparison. For this

protocol, scanner A would deliver 10*150 =

15 mGy and scanner B would deliver 8*250

= 20 mGy. In this case, even though scanner

A has a larger CTDI per 100 mAs value, the

dose to the patient needed to maintain the

same level of image quality is 30% more for

scanner B. Since it is not usually possible to

scan the same patient on each scanner and

equalize the image quality, scanning standard

phantoms with standard measurements, such

as LCD, is the only reliable way to compare

dose between two different scanner systems.

Effective Dose and DLP – From a

patient’s point of view, their only concern

is the potential health risk from the

exam. Unfortunately, there is not enough

information in the CTDI value to get an

estimate of patient risk. With CTDI there

is no indication of how much anatomy was

scanned. For the same kV, mAs, and pitch,

the CTDI value is the same whether 10 cm

of the abdomen was imaged or the entire

chest, abdomen, and pelvis. Intuitively, with

a longer scan, more of the body is exposed

to radiation and the risk should be higher.

However, CTDI does not reflect this. In order

to describe the total energy imparted to

the patient, a value called the Dose Length

Product (DLP) is used. The DLP is found

by multiplying the CTDIvol for a protocol by

the total length of anatomy scanned. The

advantage of the DLP is that it incorporates

the actual mAs used (unlike CTDI per 100

mAs) and the total scan length (unlike

CTDIvol). As such, the DLP carries much

more information about the actual dose used

than the CTDI values alone (Figure 13).

In order to estimate patient risk, it is

necessary to incorporate information about

the anatomy scanned. As previously stated,

some organs are more sensitive to radiation

than others. The radiosensitivity of each

organ is represented by a weighting factor

(w) based on the atomic bomb survivor

data3,4,5. By adding the amount of each

organ scanned, weighted by its sensitivity,

the patient’s Effective dose (E) can be

derived. One of the main benefits of E is

that it can be compared on a risk basis with

other radiation exposures. For example, in

the U.S., the average person receives about

3 millisieverts (mSv) of naturally occurring

radiation per year (mostly from radon in


CTDIW1, DLP1, E1

CTDIW2, DLP2, E2

X-rays

2x

x

E1=2 E2 if the same anatomical area is covered

Difference between these two cases:

CTDIW1 = CTDIW2 DLP1 = 2 DLP2 E1 > E2

Figure 13: Relationship between CTDI, DLP, and effective dose. For a given x-ray technique, the CTDIw value is independent of the distance covered. Therefore, the CTDIw values in cases one and two above are identical even though the patient in case one is absorbing twice as many x-rays. The DLP accounts for the distance as well as the x-ray technique and refl ects increased dose with the greater coverage. The increased DLP directly affects the effective dose which will be double in this case if the same general anatomy is covered by both scans.


the soil and cosmic rays). In comparison,

a typical CT scan of the chest is about 5-7

mSv. Therefore, the dose from the CT is

approximately equal to two years of natural

exposure. This method of describing dose

in terms of equivalent natural background

time is known as Background Equivalent

Radiation Time or BERT32. This is a useful

way of describing radiation since it places the

dose into an understandable perspective for

the patient.

Based on the the linear no-threshold model

discussed earlier, there is no “safe” level

of radiation and any exposure carries a

small but finite risk. According to the latest

Biological Effects of Ionizing radiation (BEIR)

data, the risk of developing a fatal cancer

10-20 years following an exposure of 10 mSv

is about 0.05% or 1 in 20004. To put this in

perspective, a person assumes the same risk

of death from driving 10,000 miles in a car or

from smoking a pack of cigarettes per day for

6 months14.

It is important to realize that in nearly all cases,

the benefit of medical imaging far outweighs

the small risk16, 33, 34, 35, 36. However, it is critical

that CT and other imaging modalities be used

based on accepted appropriateness criteria

and with the principles of ALARA always in

the forefront.

TOSHIBA DOSE REDUCTION

In order to adhere to the ALARA principle

and always put patient safety first, Toshiba

Medical Systems has designed its CT

scanners to give the best image quality

at the lowest possible dose. From the

dual-supported anode x-ray tube, to the

ultra-efficient Quantum detector system

and noise-free data acquisition system

(DAS), to the dose-saving SUREExposure3D

mA modulation software, to advanced,

adaptive reconstruction and noise reduction

algorithms, the Aquilion system is designed

to deliver the best image quality at the lowest

possible dose.

Quantum Detectors – In designing a dose

efficient CT scanner, the first step is to build

an acquisition system of detectors and

electronics that perform well under low signal

conditions. At the core of the Quantum

Detector is Toshiba’s patented illumination

detector material. Through a proprietary

manufacturing process, praesodynium

doped gadolinium oxy-sulfide (GOS)

crystals are sintered into a highly efficient

CT detector. The illumination material is

8 year-old pediatric lung CT.Effective dose of 1.9 mSv.

over 99% absorption efficient and optically

transparent with a light output of 2.3 times

that of cadmium tungstate (CdWO4). It is

characterized by fast decay times and low

afterglow properties that allow artifact-free

scanning down to 0.35 seconds per rotation

and below. Combined with precise and

highly shielded electronics to ensure the

quietest, truest signal possible, the Quantum

Detectors ensure maximum signal in low-

dose acquisitions.

MegaCool™ Tube – Since vibrations in the

anode can cause instabilities in the x-ray

beam requiring wider collimation with more

penumbra, Toshiba designed its MegaCool

x-ray tube with bearing supports on both

ends of the anode axis. This added anode

stability effectively reduces the variations

in the beam, allowing high-quality imaging

to be performed at fast rotation speeds,

up to 0.35 sec/rotation, while minimizing

extra patient dose from penumbra.

Furthermore, the MegaCool tube has

an innovative feature to collect off-focal

electrons and prevent them from producing

x-rays. If these electrons are not captured,

they can lead to artifacts and image quality

degradation as well as unnecessary patient

dose. By fitting a positively charged grid

near the electrically grounded anode, any

secondary, off-focal electrons are captured

and removed from the system. In this

way, the MegaCool tube provides optimum

image quality with a minimum of radiation

dose to the patient (Figure 14).

SUREExposure3D – Since the human body

is not perfectly round and uniform in size

and density, different mAs settings are

required to achieve the same image quality

in different parts of the body. For example,

during a typical lung scan, the first part of the

acquisition must penetrate the bony shoulder

area, while the middle part is largely low-

attenuating air, and the end has to penetrate

the liver and diaphragm. In this example, a

larger mAs is needed through the shoulders,

a smaller amount through the lungs, and

an amount somewhere in between for the

abdomen. Furthermore, patients are not

all the same size. A sumo wrestler would

clearly require more mAs to achieve a given

image quality level than would a ballerina.

In addition, because patients are shaped in

an elliptical fashion, more mAs is typically

needed when x-rays are passing laterally

through the body than when they are passing

anterior-posterior.


Figure 14: MegaCool x-ray tube. With its anode electrically grounded and supported at both ends, the MegaCool tube allows long scanning with no cooling delays and has a stable focal spot. This stable focal spot, along with the unique aperture that collects off-focal electrons makes the MegaCool tube an integral part of the Aquilion’s low-dose engineering.

Cathode

Aperture

Anode

Coronal MPR abdominal CT.Effective dose of 5.6 mSv.


Toshiba’s SUREExposure3D software

automatically adjusts the mAs rapidly during

the scan to adapt to and compensate for all

of these changes in attenuation level. Using

data from the AP and lateral scanograms, the

software determines exactly how much mAs

is necessary to maintain a user-defined level

of image quality (Figure 15). SUREExposure3D

does this in all three dimensions (x, y, and z).

Therefore, as the scan moves from the

shoulders to the lung, the mAs goes down,

and as the tube rotates around the patient,

less mAs is used anterior-posterior than

laterally. For the same image quality level,

compared to non-modulated scanning,

SUREExposure3D can reduce the dose by up

to 40%.

The unique nature of coronary imaging gives

another opportunity for dose savings. With

low and steady heart rates, the optimum

phase for reconstruction is typically between

65 and 80% R-R. Since the data in the rest

of the cardiac cycle is used only for examining

the bulk ventricular function, a much lower

mAs value is necessary. SUREExposure3D

with ECG dose modulation allows the mA to

be significantly reduced during the systolic

phases of the cardiac cycle, enabling a

reduction in patient dose of as much as 50%.

Boost3D – Even with an optimized detector

system and mAs modulation, highly

attenuating anatomy such as the shoulders

and pelvis can severely reduce the number

of photons reaching the detectors. This

localized reduction in photon count can

lead to degradation in image quality in the

form of excess noise and streak artifacts.

Conventionally, these highly attenuating

areas are imaged using increased mAs

and kVp to overcome the low photon

count. However, since increasing the

imaging technique results in higher patient

dose, Toshiba engineers developed an

adaptive, three-dimensional algorithm

that preferentially corrects the raw data

in areas with low photon count. This

algorithm, known as Boost3D, seeks out

portions of the raw-projection data where

there is a disproportionate loss in x-ray

signal and applies the three-dimensional

filter locally to reduce the image noise and

streak artifacts. In areas of normal signal,

no correction is applied and the native

image quality is preserved. Such local, or

adaptive, techniques produce the optimum

results because the filter is applied only

where it is needed. Since this algorithm

removes streak artifacts caused by photon

160

140

120

100

80

60

40

mA

SUREExposure

60 90 120 150 180 210 240 270 300 330

75 mAs

25 mAs

60 mAs

Figure 15: SUREExposure tailors the dose to the individual patient. More dose is needed to penetrate dense areas like the upper abdomen and bony pelvis while lower mAs is adequate in less-dense areas like the lung. Overall dose can be lowered by as much as 40% while preserving uniform image quality.

starvation, it can either be applied to

enhance images using conventional mAs

settings, or to allow low-dose imaging with

acceptable image quality by reducing

the scan technique and, thereby, the

patient dose. Figure 16A demonstrates

images through the shoulder in a cardiac

bypass patient using a relatively low scan

technique. The images exhibit typical

structured noise and streak artifacts

resulting from the low photon count.

However, when Boost3D is applied, Figure

16B shows the resulting image quality: the

image noise is greatly reduced and the

streak artifacts disappear. By reducing the

noise and mitigating the effects of low-dose

scanning, adaptive techniques such as

Boost3D are key developments in Toshiba’s

commitment to patient focused imaging.

Quantum Denoising Software

Beyond optimized scanning techniques

and streak removal, it is possible to

minimize the overall noise left in the

reconstructed image. Toshiba’s Quantum

Denoising Software (QDS) is an adaptive

noise reduction filter that works on

reconstructed image data by preferentially

smoothing areas of uniform density

while preserving the edge information

of the image. The algorithm uses locally

sampled edge information within the

image to blend together variable strength

smoothing and sharpening filters37. In

areas of uniform density with few edges,

the algorithm smoothes the image and

reduces the noise; in areas with edges,

such as near tissue boundaries and

other complex structures, the algorithm

enhances the image. QDS works in

both two and three dimensions and can

drastically reduce image noise, allowing

a corresponding savings in patient dose.

Figure 17 illustrates the substantial dose

savings possible using QDS. Figure 17A

shows the relative noise in the liver of a

patient using a standard scan technique.

17B demonstrates the increase in image

noise as the mAs is dropped by 47%.

Finally, 17C highlights the ability of the

QDS to reduce the noise in the liver to

below that of the original, higher-dose

image. QDS works in conjunction with the

SUREExposure3D software to adjust the mAs

based on the expected noise reduction

from the adaptive filter. In this way, patient

dose reduction is totally integrated in the

Aquilion console software.


Figure 16: Effect of Boost3D on streaks from low photon count. A) shows typical streaks through the shoulder region. B) shows the same dataset reconstructed with Boost3D.

Figure 17: Effect of Quantum Denoising Software on liver scan. A) shows a scan at standard x-ray exposure and noise level. B) shows a scan at 45% lower mAs. The noise value is increased. C) shows the lower mAs scan with QDS. The noise is the same as the standard exposure, but with signifi cantly less dose.

A B A B C


EMERGING TECHNOLOGIES

The next frontier of CT dose reduction

comes from emerging technologies such as

wide area coverage detectors. With the ability

to cover an entire organ, such as the heart

or the brain, in a single rotation, Toshiba’s

256 slice CT scanner (works in progress)

can drastically reduce patient dose. Without

the need for overlapping slices, wide area

detector scanners can reduce the dose by up

to 80% as compared to today’s state-of-the-

art multislice systems38. Furthermore, since

the entire heart can be acquired in a single

rotation, coronary artery imaging with doses in

the range of a calcium score is possible.

CONCLUSIONS

Comparing the patient dose between

different scanners is impossible without

also equalizing image quality. Low Contrast

Detectability is an excellent measure of the

system’s overall dose efficiency since both

dose and image quality are represented.

Many factors affect the dose, and Toshiba’s

SURETechnologies, including SUREExposure3D,

automatically adjust parameters to maintain

image quality and reduce dose. To maintain

adherence to the ALARA principal, every

aspect of Toshiba’s Aquilion CT scanner is

built to minimize patient dose.

GLOSSARY OF TERMSAbsorbed dose: The energy absorbed by

the exposed tissue or material per unit mass.

The unit is the Gray (Gy) which equals 1

Joule (energy) per kilogram.

Absorption: One possible interaction of

radiation with matter. In absorption, the

particle (x-ray, electron, etc) is completely

stopped and deposits all of its energy in the

tissue or material.

ALARA: As Low As Reasonably Achievable.

Principle of radiation protection that

requires the use of the minimum amount

of radiation to achieve the clinical task.

Not necessarily the “lowest” dose, but the

lowest appropriate dose.

Attenuation: The reduction in radiation from

passing through a tissue or material. This

takes place through absorption or scatter of

the x-ray photons.

Beam hardening: The preferential removal

of low- and medium-energy x-rays from

the spectrum by absorption in a material.

The material acts as an x-ray filter. If these

changes in the x-ray spectrum are not

accounted for in the reconstruction, beam

hardening artifact can occur which appears

as streaks and shadows in the CT image,

especially near very dense objects.

Beam Pitch: Pitch value based on the entire

x-ray beam width. Often just referred to as

“pitch”. Identical to IEC definition of pitch. It

is defined as the table feed divided by the

nominal slice thickness times the number

of slices. Along with the rotation time,

Beam Pitch determines the speed at which

the volume is scanned. It is different from

Helical Pitch in that it divides out the total

number of slices. Typical values are between

0.5 and 1.5. Beam Pitch values below 1

indicate an overlapping exposure.

Body phantom: The standard 32 cm

diameter cylindrical phantom used to

measure CTDI. It is intended to represent

the attenuation and scatter properties of a

typical abdomen. The phantom is made of

Perspex (PMMA), is approximately 15 cm

deep, and typically has 5 holes (one hole

at the center and four holes 1 cm below

the surface at the 12, 3, 6, and 9 o’clock

positions) to hold the 10 cm ion chamber.

Bowtie Filter: A beam shaping filter that is

thicker at the edges than in the center. It

is used to reduce the beam intensity at the

edges of the field of view where the patients

are not as thick. This both reduces the

patient dose and makes for a more uniform

photon flux at the detectors.

Cadmium Tungstate (CdWO4): High-density

crystal material used in gamma- and x-ray

detection. Most of the original solid state CT

detectors were CdWO4 and its light output

value is the standard against which other

materials are compared. Most manufacturers

have moved away from CdWO4 due to the

toxic nature of cadmium.

Cone Angle: Angle at which the x-ray beam

diverges from the tube in the z-direction.

MSCT scanners with 16 slices and greater

have cone angle that is large enough that it

must be accounted for in reconstruction to

avoid image artifacts. The most successful

cone beam reconstruction algorithms in use

in scanners up to 64 slices are based on the

FDK (Feldkamp Davis Kress) algorithm.

CTDI: Computed Tomography Dose Index.

The standard method of measuring an index

for dose in a CT system. While this number

is an index and does not represent an actual

patient dose, it is benchmark. It is measured

using two Perspex (PMMA) phantoms: a 32

cm diameter “body” phantom and a 16 cm

diameter “head” phantom. CTDI is a useful

metric to compare the doses from different

protocols on the same scanner. However,

since it is so dependent on filtration, it is

unreliable to use to compare the doses from

different scanners. CTDI is not proportional to

patient dose since its value is the same whether

10 cm or 100 cm of the body is exposed.

CTDI100: The CTDI value measured using

a 100 mm ion chamber. CTDI¬100 is the

integrated absorbed dose value along the

entire 100 mm chamber. As scanners

increase z-axis coverage, the 100 mm

chamber becomes increasingly inadequate

to evaluate the dose.

CTDIFDA

: The CTDI value measured by

integrating 14 slice widths. This value is

difficult to measure since the integration

length is variable.

CTDIvol: The weighted CTDI (CTDIw)

divided by the helical pitch. By correcting

for the pitch, this value represents the

overlapping dose profiles for pitches < 1

or the space between the dose profiles for

pitches >1.

CTDIw: Weighted CTDI. The weighted

average of the CTDI100 values calculated

by adding 2/3 of the average measured

peripheral value to 1/3 of the measured

center value.

Dose Profile: The actual collimated width

of the exposure beam. For MSCT, the

dose profile is usually slightly larger than

the nominal collimation to ensure uniform

radiation density in all the detector rows.

DLP: Dose Length Product. The DLP

represents the total amount of energy

imparted to the body. It is the product

of the CTDIvol and the scan length. DLP

is useful when comparing the dose

from different scanners since it contains

information on the scan protocol used. DLP

is proportional to patient dose. It is defined

in units of mGy-cm.

Effective dose (E): A value for dose that

represents patient risk in terms of an

equivalent whole body exposure. It allows

for dose comparisons with other sources

of ionizing radiation including different

medical imaging modalities and natural

background radiation. The risks estimates

are largely based on studies of the

atomic bomb survivors. Effective dose is

expressed in units of milliSieverts (mSv).

Fan angle: Angle at which the x-ray beam

diverges from the tube in the xy-direction.

This fan angle allows the tube and detectors

to rotate around the scanner’s isocenter and

image circular fields of view (FOVs) up to 50

cm and more.

Field of View (FOV): The diameter of the

reconstructed image. The FOV determines

how much of the anatomy is visible in the

image. It is possible to reconstruct images

up to the acquired FOV.

Filtration: Thickness of aluminum or copper

(typically) used to remove low energy x-rays

from the beam. These low energy x-rays

only add to patient dose and do not take

part in the creation of images. The amount

of filtration also affects the total tube output.

GOS: Gadolinium Oxysulfide (Gd2O2S).

Dense scintillator material used to make

fast CT detectors with high light output.

preferably doped with Praesodynium (Pr)

to improve the light output.

Gray (Gy): Unit of absorbed dose. The

Gy is the SI unit equivalent of the rad (1

Gy = 100 rad). Absorbed dose is defined

in terms of energy absorbed in a unit of

mass. The Gy is equal to 1 Joule of energy

absorbed in 1 kg of tissue.

Head phantom: The standard 16 cm

diameter cylindrical phantom used to

measure CTDI. It is intended to represent

the attenuation of a typical head. The

phantom is made of Perspex (PMMA), is

approximately 15 cm deep, and typically

has 5 holes (one hole at the center and four

holes 1 cm below the surface at the 12, 3,

6, and 9 o’clock positions) to hold the 10 cm

ion chamber.

Helical Pitch: Pitch value based on the

width of the individual slice in a multislice

scanner. It is defined as the table feed

divided by the nominal slice. Along with

the rotation time, Helical Pitch determines

the speed at which the volume is scanned.

It differs from the Beam Pitch in that it is

not divided by the total number of slices.

Typical values of helical pitch for a 64 slice

system are in the range of 41 to 96. Typical

values of helical pitch for a 16 slice system

are in the range of 11 to 24.

Hounsfield Unit: Named after Sir Godfrey

Hounsfield who defined it. The HU

represents the attenuation properties of the

tissue. It is also referred to as a CT number.

It is equal to the (µx – µw / µw) x 1000 such

that air is -1000 and water is 0.

Ionization Chamber: A radiation measuring

device used to estimate dose in a CT system.

It detects the x-rays by counting the number

of charged particles (ions) created by the

exposure. It is sometimes called an ion

chamber, for short. Typically, a 100 mm ion

chamber is used to measure CTDI.

Kerma: The sum of the energy of all the

charged particles (ions) created by the

exposure. For low energies, kerma is

roughly equal to absorbed dose. For high



energies, the kerma can be much higher

than the absorbed dose.

K-factor: Conversion factor for estimating

effective dose from DLP39. There are

different values for the head, neck, chest,

abdomen, and pelvis. This is a convenient

way to get a rough estimate of the effective

dose for a patient using values readily

available on the scanner.

kV: The energy between the x-ray tube’s

anode and cathode. This determines

the maximum energy of the x-ray beam

produced by the tube and also influences the

average energy of the spectrum. Higher kV

values lead to higher energy x-rays which will

penetrate thick objects more easily. However,

higher energy also means higher dose.

Low Contrast Detectability (LCD): The

system’s ability to depict small objects with

densities very near that of their surrounding

tissue. Measurements of LCD include

the detectable object size, the percentage

difference from the background, and the

dose needed to achieve the object size

and contrast. Since the LCD is an image

quality metric that is sensitive to dose, it is

an excellent measurement of a CT system’s

overall dose efficiency.

mA: Tube current. The mA is the the amount

of current placed through the x-ray tube’s

cathode. The current is proportional to the

number of electrons created at the cathode

and accelerated towards the anode by the

high voltage (kV). More mA leads to more

electrons which leads to more x-rays leaving

the tube. The dose delivered by the system

varies linearly with mA.

mAs: The mAs (pronounced “mass”) is the

product of the tube current and the rotation

time. This gives the total number of x-rays

produced over a given time. Dose varies

linearly with mAs. The mAs contains more

information than the mA alone (300 mA at 1

sec rotation is double the dose of 300 mA at

0.5 sec rotation).

mAseff

: Effective mAs. The mAseff

incorporates the overlapping nature of helical

scanning by dividing the mAs by the beam

pitch. By including the extra dose from the

overlapping beams, mAseff

contains more

information than the mAs alone (100 mAs

at a pitch of 0.5 has double the dose of 100

mAs at pitch 1).

MPR: Multiplanar Reformat. Method of

taking a volume of thin slice data and

resampling it in a different orientation than

the data was acquired. With the isotropic

resolution of MSCT scanners, MPRs are

routinely created with no difference in

spatial resolution than the original axial

dataset. Since the anatomy of interest is

not usually best visualized in the axial plane,

the MRP is an invaluable tool for analyzing a

MSCT dataset.

MSAD: Multiple Scan Average Dose. The

average dose in the central slice with multiple

adjacent slices being exposed. It is equal

to the CTDI when the slice spacing is 1 and

enough scans are acquired.

MTF: Modulation Transfer Function.

Describes the scanner’s ability to represent

different spatial frequencies. In other

words, how faithfully a scanner can resolve

the details of the object scanned. It is

measured by taking the Fourier transform

of the point spread function of the scanner

and can be used to describe the scanner’s

spatial resolution.

Noise: Random variation in the pixel values

of the reconstructed image. Noise contains

no useful information and degrades image

quality. High noise values can obscure

anatomy and hinder low contrast detectability.

On a given scanner, noise increases with

decreasing dose.

Noise power spectrum (NPS): A more

complete description of the image noise than

simple standard deviation alone. NPS reveals

correlations in the noise by representing noise

in terms of its spatial frequency distribution.

Organ dose: Absorbed dose to a given organ.

This value is particularly useful in formulating

a true effective patient dose.

Partial volume artifact: Image artifact

occurring when the pixel size or slice

thickness is not small enough to individually

represent multiple densities. Usually appears

as streaking, blurring, or loss of

edge definition.

Pitch: See Beam Pitch and Helical Pitch.

Quantum Detectors: Toshiba’s highly efficient

detection system consisting of the patented,

GOS-based illumination™ detector material

and the quiet, efficient Data Acquisition

System (DAS). Toshiba’s Quantum detectors

provide 0.5 mm slices and 2 mm low-contrast

resolution resulting in excellent image quality.

Rad: Unit of absorbed dose (stands for

Radiation Absorbed Dose) equal to one erg

absorbed in one gram of tissue. The rad has

been replaced by the SI unit Gray (Gy). One

Gy equals 100 rad.

Rem: Unit of effective dose (stands for

Radiation Equivalent Man). The rem has

been replaced by the SI unit Sievert (Sv).

One Sv equals 100 rem.

Scatter: X-rays that are not completely

absorbed or completely transmitted will

deposit a portion of their energy and continue

on in a different direction. This scattered

radiation contributes to noise within the

image when they are picked up at the

detector.

Sievert (Sv): Unit of effective dose. The

Sievert replaces the rem where 1 Sv is

equal to 100 rem. The average natural

background radiation in the United States is

about 3 millisSeverts (mSv).

Slice sensitivity profile: The full width at half

maximum (FWHM) size of the reconstructed

slice in the z-direction. Factors such as

the detector width, helical pitch, and z-

interpolation algorithm can affect the slice

sensitivity profile.

SUREExposure3D: Toshiba’s automatic

exposure control and mA modulation

software. SUREExposure3D controls

and modulates the mA in the XY and

Z directions to achieve and maintain a

uniform, user-selected noise level within the

images. SUREExposure3D uses attenuation

information gathered from the AP and lateral

scanograms to determine the appropriate

mA. SUREExposure3D can lower the dose by

as much as 40% while maintaining image

quality. Since SUREExposure3D is based on

the actual patient attenuation, it is especially

useful for ensuring that the appropriate dose

levels are used in pediatric cases.

Table Feed: The amount of table travel (in

mm) per rotation, for helical scanning, or

between scans for axial scanning. Along

with the detector collimation, table feed

determines the pitch.

Voxel: The discrete image volume is made

of individual 3D elements. Each element is

called a voxel (short for volume element) and

represents a single HU value.

REFERENCES1. Amis Jr ES, Butler PF, Applegate KE,

Birnbaum SB, Brateman LF, Hevezi JM,

Mettler FA, Morin RL, Pentacost MJ, Smith

GG, Strauss KJ, Zeman RK. 2007, American

College of Radiology White Paper on

Radiation Dose in Medicine, J Am Coll Radiol

4:272-284.

2. Medical Imaging Technology Alliance.

2007, How Innovations in Medical Imaging

Have Reduced Radiation Dosage. NEMA,

Rosslyn, VA, Available Online: http://www.

medicalimaging.org/news/fullreport_reduced_

radiation_dose.pdf.

3. Nuclear Regulatory Commission. 1990,

Health effects of exposure to low levels of

ionizing radiation: BEIR V. Washington, DC:

National Academy Press.

4. Nuclear Regulatory Commission. 2006,

Health Risks from Exposure to Low Levels

of Ionizing Radiation: BEIR VII Phase 2.

Washington, DC: National Academy Press.

5. International Commission on Radiological

Protection. 1990, Recommendations of the

International Commission on Radiological

Protection. Publication 60, International

Commission on Radiological Protection.

Oxford, England: Pergamon.

6. Thompson DE, Mabuchi K, Ron E, et al.

1994, Cancer incidence in atomic bomb

survivors. Part II. Solid tumors, 1958-1987.

Radiation Research 137(suppl 2):S17-S67.

7. Pierce DA, Preston DL. 2000, Radiation-

related cancer risks at low doses among

atomic bomb survivors. Radiation Research,

154(2):178-86.

8. Brant-Zawadzki M. 2005, Diagnostic

Radiology: Major Weapon in Patient Care

or Weapon of Mass Destruction? J Am Coll

Radiol. 2(4):301-303.

9. Ron E. 2003, Cancer risks from medical

radiation. Health Phys. 5(1):47-59.

10. Valentin J. 2005, Low-dose extrapolation

of radiationrelated cancer risk. Ann ICRP.

35(4):1-140.

11. Upton AC, Adelstein SJ, Brenner DJ,

et al. 2001, Report No. 136-Evaluation of

the Linear-Nonthreshold Dose-Response

Model for Ionizing Radiation. Bethesda, MD:

National Council on Radiation Protection and

Measurements (NCRP)

12. Nations Scientific Committee on the

Effects of Atomic Radiation UNSCEAR.

2000,Sources and Effects of Ionizing

Radiation: United Report to the General

Assembly, With Scientific Annexes. New York,

NY: United Nations.

13. Cox R, Muirhead CR, Stather JW,

Edwards AA, Little MP. 1995, Risk of

radiation-induced cancer at low doses

and low dose rates for radiation protection

purposes. Documents of the NRPB. 6:1-77.

14. Hall EJ. 1994, Radiobiology for the

Radiologist. Fourth Edition. J.B. Lippincott

Company.

15. Cember H. 1996, Introduction to Health

Physics. 3rd ed. McGraw-Hill. pg 290

16. Frush DP, Slack CC, Hollingsworth CL,

et al. 2002, Computer-simulated radiation

dose reduction for pediatric abdominal

multidetector CT. AJR Am J Roentgenol.

179:1107-1113



17. Brenner DJ, Elliston CD, Hall EF, Berdon

WE, 2001, Estimated risks of radiation

induced fatal cancer from pediatric CT. Am J

Roentgenol; 176(2):289-296.

18. Paterson A, Frush DP, Donnelly LF. 2001,

Helical CT of the body: are settings adjusted

for pediatric patients? AJR 176:297-301

19. Cohen G, Wagner LK, Amtey SR, Di

Bianca FA. 1981, Contrast-detail-Dose and

dose efficiency analysis of a scanning digital

and a screen-film-grid radiographic system.

Med. Phys. 8, 358

20. AAPM Report 39: Specification and

Acceptance Testing of Computed Tomography

Scanners. May 1993

21. Ishida T, Tsukagoshi S, Kondo K, Kainuma

K, Okumura M, Sasaki T. 2004, Evaluation of

dose efficiency index compared to receiver

operating characteristics for assessing CT low-

contrast performance. Proceedings of SPIE

-- Volume 5368 Medical Imaging, Physics of

Medical Imaging, Martin J. Yaffe, Michael J.

Flynn, Editors, pp. 527-533

22. Johns HE, Cunningham JR. 1983, Ch

XVI: Diagnostic Radiology. In The Physics of

Radiology, Fourth Edition, Thomas, p624-

625.

23. Bushberg JT, Seibert JA, Leidholdt EM,

Boone III, JM. 2002, Ch 13 Computed

Tomography pp. 367-369. In The Essential

Physics of Medical Imaging, Second Edition.

Lipincott Williams and Wilkins.

24. Seeram E. 2001, Computed Tomography:

Physical Principles, Clinical Applications, and

Quality Control. Second Edition. Saunders.

25. McNitt-Gray MF. 2002, Radiation dose in

CT. RadioGraphics 22:1541-1553.

26. Huda W. 2002, Dose and image quality in

CT. Pediatr Radiol 32(10):709-713.

27. Okumura M, Tamatani M, Igarishi K.

2002, Development of X-ray Detector for

Multi-slice CT with 0.5 mm. Slice Thickness

and 0.5 Second Revolution. Proceedings of

the SPIE Vol. 4682.

28. Huda W, Lieberman KA, Chang J,

Roskopf ML. 2004, Patient size and x-

ray technique factors in head computed

tomography examinations. I. Radiation doses.

Med. Phys. 31 (3):588-94.

29. DeMarco JJ, Cagnon CH, Cody DD,

Stevens DM, McCollough CH, Zankl M,

Angel E, McNitt-Gray MF. 2007, Estimating

radiation doses from multidetector CT using

Monte Carlo simulations: effects of different

size voxelized patient models on magnitudes

of organ and effective dose. Phys Med Biol.

52(9): 2583-97.

30. Jangland L, Sanner E, Persliden J. 2004,

Dose reduction in computed tomography by

individualized scan protocols. Acta Radiol.

45(3):301-7.

31. Shope TB, Gagne RM, Johnson GC.

1981, A method for describing the doses

delivered by transmission x-ray computed

tomography. Med Phys. 8(4):488-95.

32. Cameron JR. 1991, A radiation unit for the

public. Physics and Society News 20:2.

33. Coren ME, Ng V, Rubens M, Rosenthal

M, Bush A. 1998, The value of ultrafast

computed tomography in the investigation

of pediatric chest disease. Pediatr Pulmonol

26:389-395.

34. Mettler FA Jr, Wiest PW, Locken JA, et al.

2000, CT scanning: patterns of use and dose.

J Radiol Prot29:353 -359.35.

35. Haaga JR. 2001, Radiation dose

management: weighing risk versus benefit.

AJR 177:289-291.

36. Huda W. 2007, Radiation Doses and

Risks in Chest Computed Tomography

Examinations. ProcAmThoracSoc, 4: 316-

320.

37. Okumura M, Toyoshima N, Katada K,

Sasaki M. 2003, Performance Evaluation for

Image Filters Used for CT Images. Abstract of

RSNA, p720.

38. Mori S, Endo M, Nishizawa K, Murase K,

Fujiwara H, Tanada S. 2006, Comparison of

patient doses in 256-slice CT and 16-slice CT

scanners. Br J Radiol, 79:56-61.

39. European guidelines for quality criteria for

computed tomography. 2000, Luxembourg:

European Commission.