ct from past to future carlo maccia medical physicist caats 43 bd du maréchal joffre –...

CT from past to future

Carlo MacciaMedical Physicist

CAATS 43 Bd du Maréchal Joffre – Bourg-La-Reine – FRANCE

XI. National Turkish Medical Physics Congress 14-18 November 2007 - Antalya

Content

CT equipment and technology

Recall of basic physical principles of CT

Radiation protection rules and QC

CT dosimetry quantities

Reference Dose values and Quality criteria for CT images

INTRODUCTION

Computed tomography (CT) was commercially introduced into radiology in 1972 and was the first fully digital imaging device making it truly revolutionary in diagnostic imaging. In 1979, Godfrey Hounsfield and Allen Cormack were awarded the Nobel Prize in Physiology and Medicine for their contributions in the development of CT.

CT differs from conventional projection imaging in two significant ways:• CT forms a cross-sectional tomographic image, eliminating the

superimposition of structures that occur in plane film imaging because of the compression of three-dimensional body structures into the two-dimensional recording system

• the sensitivity of CT to subtle differences in x-ray attenuation is at least a factor of 10 higher than normally achieved by film-screen recording systems

THE BASIC PHYSICS PROBLEM

Under ideal conditions (monochromatic beam, ideal collimation, perfect detection, etc)x-ray intensity observes an exponential decay law:

N = N0 e-x

where N0 and N are the intensities of the incident and exiting x-rays, respectively, xis the path length through the attenuating material, and is the linear attenuationcoefficient of the material along the path x.

ASIDE

If we had a block consisting of a single attenuating material with unknown , we

could measure its length (x) and the incident (N0) and exiting intensities (N) , and

then solve for .

Now suppose we have an object with unknown contents, we can make a measurement of x-ray attenuation along a straight line through it

but for all intense of purposes all that this will tell us is a single number representing

the total attenuation of the material in the path. What we really want is the

attenuation coefficient at each position along the path.

So essentially we have

and thus

With a single transmission measurement, the separate attenuation coefficients cannot

be determined because there are too many unknown values of i where i = 1,2,3 ,…, n.

In order to solve this equation for the n values of i we will need n2 independent

transmission equations (the above equation would be one of the nn22 required equations).Consider the case for n = 4 and each block had a size x:

We can see from the above illustration that in order to solve for 1, 2, 3 and 4, we would need 4 independent equations (N1, N2, N3 and N4).

DATA ACQUISITION GEOMETRIES

A variety of geometry's have been developed to acquire the x-ray transmission data needed for image reconstruction in CT. Some geometry's have been tagged as a “generation” of CT scanner and these labels are useful in different scanner designs.The following scanner geometry's, data acquisition modes and primary technologies have been used to date:

• First Generation CT Scanner (EMI, 1973)• Second Generation CT Scanner (1974) • Third Generation CT Scanner (GE & Siemens, 1975-76)• Fourth Generation CT Scanner (1977)• Low Voltage Slip Ring Technology (Siemens, 1982)• Fifth Generation CT Scanner (1984)• Spiral CT Scanner (Siemens, 1988)• Multi-slice CT Scanner (Dual-slice, Elscint, 1992)• Multi-slice CT Scanner (Quad-slice, 1998)• Dual source CT (64-slice with two X-ray tubes, Philips 2006, 256-slice

Toshiba)

FIRST GENERATION CT SCANNER (Translate/Rotate)

First Head Scanner

NOTENOTE• this method was theoretically immune to the effects of scattered x-ray (single detector system)• because of the long scan times, this method of scanning was applicable to scanning of parts of anatomy that could have been kept motionless, such as the head

THIRD GENERATION CT SCANNER (Rotate/Rotate)

Predominant design of current commercially

available CT scanners

LOW VOLTAGE SLIP RING TECHNOLOGY

Some third and fourth generation CT scanners employ a slip ring to supply power andreceive signals from rotating parts. In the slip-ring method, an electrical conductivebrush moves along a ring-shaped electrically conductive rail. The use of a slip ringpermits high-speed continuous scanning, and dramatically increases both the performance and range of clinical applications of CT scanning.

• allows for 1 second ( or < 1 second or sub-second) scan times• allows for helical (or volumetric) scanning

A look inside a rotate/rotate CTA look inside a rotate/rotate CT

X-Ray Tube

Detector Arrayand Collimator

A Look Inside a Slip Ring CT

X-RayTube

Detector Array

Slip Ring

Note: how most

of theelectronics

isplaced on

the rotatinggantry

SPIRAL CT SCANNERS (Conventional Scanning Mode)

SPIRAL CT SCANNERS (Helical Scanning Mode)

• If the x-ray tube can rotate constantly, the patient can then be moved continuously through the beam, making the examination much faster

MULTI-SLICE CT SCANNERS (Dual Slice)

MULTI-SLICE CT SCANNERS (Quad Slice)

To build quad-slice spiral CT scanners, manufacturers had to develop detector arcs with more than four elements in the longitudinal (z) axis direction, creating a curved two-dimensional detector arrays.

GE Scanners

Dual Source CTSingle source CT

Fast + Poor Image quality Fast + Improved Image quality

AXIAL IMAGE RECONSTRUCTION

The task of reconstruction is to compute an attenuation coefficient for each picture element (pixel) and then to assign a CT number to each of these elements.

• in order to create multiple projections in a single 360° tube rotation, during a single projection the x-ray tube is pulsed and the detector array is sampled after each pulse

IMAGE RAY SUM

A B

• For a single detector, a ray sum consists of all the linear attenuation coefficient data along the corresponding x-ray beam path (eg: path AB)• For a single x-ray beam path, the ray sum is not the simple summation of the attenuation coefficients of the intercepted pixels.

Recall from previous lecture notes

Pixel Position Output Intensity

I1 = I0 e-1w1

I2 = I1 e-2w = I0e-w(1 + 2)2

n In = I0e-w(1 + 2 + … + n)

therefore

I0 In

1 + 2 + 3 + … + n = 1 ln (I0/ In) w

Ray Sum Value

Actually, the ray sum value that is computed is proportional to the sum of the n attenuation coefficients along the x-ray beam path

IMAGE PROJECTION

• typically anywhere between 800 - 1000800 - 1000 projections are collected in oneone 360° tube rotation to reconstruct a singlesingle axial image

• a projection is defined as the set of ray sums measured in all detectors during a single x-ray tube pulse

Detector Position

Images slices are reconstructed into a matrix consisting of multiple volume elements (voxel) each with a unique value.

PROBLEMThe volume scanned in a single rotation differs between the conventional and helical scanning methods.

IMAGE INTERPOLATION (SPIRAL CT)

ANSWERInterpolate desired axial image from volume data set prior to image reconstruction.

VOLUME ELEMENT (VOXEL)

New CT Features

• The new helical scanning CT units allow a range of new features, such as : CT fluoroscopy, where the patient is stationary,

but the tube continues to rotate multislice CT, where up to 64 (128 - 256) slices

can be collected simultaneously 3-dimensional CT and CT endoscopy Cardiac image acquisition during relevant heart

phases (ECG pulsing synchronization)

• Real Time Guidance (up to 8 fps)

• Great Image Quality• Low Risk• Faster Procedures

(up to 66% fasterthan non fluoroscopicprocedures)

• Approx. 80 kVp, 30 mA

CT Fluoroscopy

Content






The final result of the CT image reconstruction is an accurate estimate of the x-rayabsorption values characteristic of individual voxels.

CT Number = 1000 p - w = Hounsfield Unit (HU) w

where p is the linear attenuation value assigned to a given pixel and w is the linearattenuation value of water.

ASIDE• w is obtained during calibration of the CT scanner• by definition, the HU of water is 0 and the HU value for air is -1,000• above equation defines 100 HU as equal to a 10 % difference in the linear attenuation coefficient relative to water• the value 1000 in the numerator is a scale factor and determines the contrast scale

CT NUMBER

FIELD-OF-VIEW (FOV)

FOV is the diameter of the area being imaged (e.g: 25 cm Head and 35 cm Body scan)

• CT pixel size is determined by dividing the FOV by the matrix size (typically 512 x 512 – 768 x 768 or 1024 x 1024)

IMAGE DISPLAY• reconstructed images are viewed on a CRT monitor or printed onto film using a

laser printer• each pixel is normally represented by 12 bits, or 4096 gray levels, which is larger

than the display range of monitors or film• window width and level are used to optimize the appearance of CT images by determining the contrast and brightness levels assigned to the CT image data

IMAGE QUALITY

Image quality may be characterized in terms of:

• contrast• noise• spatial resolution

ASIDE• in general, image quality involves tradeoffs between these three factors and patient dose. • artifacts encountered during CT scanning can degrade image quality

IMAGE CONTRAST

CT contrast is the difference in the HU values between tissues. This contrast generallyincreases as kVp decreases but is not affected by mAs or scan time.

• CT contrast may be artificially increased by adding a contrast medium such as iodine• image noise may prevent detection of low-contrast objects such as tumors with a density close to the adjacent tissue• the displayed image contrast is primarily determined by the CT window width and window level settings.

CT Photon Energy Range (120 or 140 kVp)

LOW CONTRAST RESOLUTION

• Measurement TechniqueCatphan 500 (phantom)

Insert Diametre : 2 mm to 15 mm.

Contrast levels : 0.3, 0.5 and 1%

Supra slice (Periphery)

Z = 40 mm

Subs slice (centre)

Z= 3, 5, 7 mm

1%0,5%

0,3%

IMAGE NOISE

The sources of image noise in CT are:

• quantum mottle (the number of photons used to make an image)• inaccuracies in the image reconstruction process (software filter phase); and• electronic noise introduced after detection

Noise in CT is usually defined as the standard deviation () of the CT numberscalculated from pixel values in a predefined region-of-interest (ROI) using an imageof a uniform material (usually water). The selected ROI region should be void ofobjects and cover a sufficiently large image area (circular diameter > 10 mm).

For GE scanners:

ROI CT number Average Value = 0.0 3.0 HU ROI CT number Standard Deviation = 3.5 0.7 HU

ROI Area = 13.17 cm2

Mean = 1.75 Std. Dev. = 2.9

Scan Parameters: Small Scan FOV, 25 cm DFOV, 5122 Matrix, Standard Resolution, Peristaltic Option OFF, 13.17 cm2 CROI, Normal Scan Type, 5 mm slice thickness, 170 mA and 2 sec scan time

NOTE: Noise = 2.9

ELECTRONIC NOISE

• in modern CT scanners electronic noise is kept to a minimum• a CT scanner whose noise is dominated by the detection of a finite number of x-rays

(quantum mottle) is called quantum limited• in a quantum limited CT scanner

(noise)2 1 patient dose

• a CT scanner can be shown to be quantum limited by plotting

(noise)2 vs 1 (any parameter that affects patient dose)

and determining the magnitude of the y-intercept of the interpolated linear curve fit

Since in a quantum limited CT scanner

(noise)2 1 patient dose/pixelthen

(noise)2 1 B • D • H • w3

where B - is the fractional transmission of the patient D - is the maximum surface dose ( mAs) H - is the slice thickness w - is the reconstructed pixel width

• quantum mottle (and thus noise) decreases as the number of photons increases• CT noise is generally reduced by increasing the kVp, mA or scan time (if all

other parameters are kept constant)• CT noise is also reduced by increasing voxel size (ie: by decreasing matrix

size, increasing FOV or increasing the slice thickness)• typically noise with a modern CT scanner system is approximately 5 HU (or 0.5% difference in attenuation coefficient)

IMAGE RESOLUTION

Spatial resolution is the ability to discriminate between adjacent objects and is afunction of pixel size.

• If the CT FOV is D and the matrix size is M, then pixel size is D/M. Example:• For a typical head scan with a FOV of 25 cm and a matrix of 512 pixels, the pixel

size is 0.5 mm• Because two pixels are required to define a line pair (lp), the best achievable

spatial resolution is 1 lp/mm

• typically resolution in CT scanning ranges from 0.5 to 1.5 lp/mm• the axial resolution may be improved by operating in a high resolution mode

using a smaller FOV or a larger matrix size• factors that may also improve CT spatial resolution by reducing image blur include smaller focal spots, smaller detectors and more projections• resolution perpendicular to the section is dependent on slice thickness and is

important in Sagittal and Coronal image reconstruction

IMAGE RESOLUTION

• Measurement Technique

• MTF (Modulation Transfer Function) objective method

• Assessment of a bar pattern – subjective method

IMAGE RESOLUTION

• MTF can be considered as a reliable measure of the information transfer from the object to the image. It illustrates, for each individual spatial frequency, the progressive degradation of the signal due to the system in terms of % of contrast loss.

IMAGE RESOLUTION

• The MTF is assessed from the Fourier Transform of the Linear Spread Function (LSF) which is a measure of the ability of a system to form sharp images; it is determined by measuring the spatial density distribution on film of the X-ray image of a narrow slit in a dense metal, such as lead.

• The point spread function (PSF) describes the response of an imaging system to a point source or point object

IMAGE RESOLUTION

The image of the « point object » is not a single point but a set of different points representing the degradation of the signal.

IMAGE RESOLUTION

• MTF curves at 50 %, 10 % and 2 %.

PQ 5000

IMAGE RESOLUTION

Typical values• Standard mode : 7 line pairs / cm .• Maximum values : 17 to 18 line pairs / cm (high resolution

mode)

IMAGE RESOLUTION (influencing factors)

• Acquisition• Number of projections


• Acquisition• Number of projections (floating focal spot

technique)


• Acquisition• Actual detector aperture• The smaller detector aperture the better spatial

resolution• Slice thickness (reduction of scattered radiation,

improvement of image sharpness)

IMAGE RESOLUTION (Z-Axis)

• Z-axis resolution is important for 3D reconstruction ==> Isotropic dimension of the pixel

• Z-axis resolution – Slice thickness

– Pitch

Abdomen, Pelvis

Abdomen, Pelvis ChestChest AngiographyAngiography

IMAGE RESOLUTION (Z-Axis)

• If, within the slice, the object shows a continuity along the Z-axis, the HU remain constant

• If, within the slice, the object is not continuous, the partial volume effect would change the HU value

SLICE THICKNESS

• Measured at the isocentre of rotation• Allow to check the overlapping of adjacent slices• Expressed in terms of image profile at the Full

Width at Half Maximum (FWHM) value

Note : Θ = 45° magnification factor = 1 Θ = 63.5° magnification factor = 2

Content






SLICE THICKNESS

• Catphan 500 Phantom• Θ = 23°

DOSIMETRIC QUANTITIES C.T.

• CTDI (Computed Tomography Dose Index)

• DLP (Dose-Length Product)

• MSAD (Multiple Scan Average Dose)

The CTDI is the integral along a line parallel to the axis of rotation (z) of the dose profile (D(z)) for a single slice, divided by the nominal slice thickness T

In practice, a convenient assessment of CTDI can be made using a pencil ionization chamber with an active length of 100 mm so as to provide a measurement of CTDI100 expressed in terms of absorbed dose to air

(mGy).

D(z)dz T

1 =

+

-

CTDI

COMPUTED TOMOGRAPHY DOSE INDEX (CTDI)


• Measurement principle

e

Aire= e x CTDI

CTDI Airee

e n x e Z mm

Mean Dose.

Ionization Chamber

• measurements of CTDI may be carried out free-in-air in parallel with the axis of rotation of the scanner (CTDI100, air)

• or at the centre (CTDI100, c) • and 10 mm below the surface

(CTDI100, p) of standard CT dosimetry phantoms.

• the subscript `n' (nCTDI) is used to denote when these measurements have been normalised to unit mAs.


Air

1 cm

Centre

Ideal

HETEROGENEITY OF DOSE PROFILES

On the assumption that dose in a particular phantom decreases linearly with radial position from the surface to the centre, then the normalised average dose to the slice is approximated by the (normalised) weighted CTDI: [mGy(mAs)-1]

where:– C is the tube current x the exposure time (mAs)

– CTDI100,p represents an average of measurements at four

different locations around the periphery of the phantom

)( CTDI3

2 + CTDI

3

1

C

1 = CTDI p100,c100,wn


Two reference dose quantities are proposed for CT in order to promote the use of good technique:

– CTDIw in the standard head or body CT dosimetry phantom for a single slice in serial scanning or per rotation in helical scanning : [mGy]

where:– nCTDIw is the normalised weighted CTDI in the head or body phantom

for the settings of nominal slice thickness and applied potential used for an examination

– C is the tube current x the exposure time (mAs) for a single slice in serial scanning or per rotation in helical scanning.

C CTDI = CTDI wnw

REFERENCE DOSE QUANTITIES

• CTDI(vol) for non adjacent slices : [mGy]

• Axial mode CTDI(vol) = CTDI(w) x T Slice interspace

• Helical Mode CTDI(vol) = CTDI(w)

Pitch


• DLP Dose-length product for a complete examination : [mGy • cm]

where :– i represents each serial scan sequence forming part of

an examination

– N is the number of slices, each of thickness T (cm) and radiographic exposure C (mAs), in a particular sequence.

N.B.: Any variations in applied potential setting during the examination will require corresponding changes in the value of nCTDIw used.

iC N T CTDI = DLP wn


In the case of helical (spiral) scanning [mGy • cm] :

where, for each of i helical sequences forming part of an examination :

– T is the nominal irradiated slice thickness (cm)

– A is the tube current (mA)

– t is the total acquisition time (s) for the sequence.

N.B. : nCTDIw is determined for a single slice as in serial scanning.

t A T CTDI = DLP wni


• Multiple Scan Average Dose (MSAD) : The average dose across the central slice from a series of N slices (each of thickness T) when there is a constant increment I between successive slices:

where:

DN,I(z) is the multiple scan dose profile along a line parallel to the axis of rotation (z).

(z)dz D = MSAD IN,2I

+

2I

-

1Ι


Multiple Scan Average Dose

(z)dz D = MSAD IN,2I

+

2I

-

1Ι

e

e

Z mmT T

Pitch =1 ; CTDI=MSAD I

MSAD : dose delivered while scanning with non adjacent slices (axial mode)

• Scanned area Larger collimation ==> 40 mm« Important irradiated volume : overscan »

• Speed Rotation time 0.33 to 0.5 s – Matrix Size 512 x 512 to 1024 x 1024 (Philips).

• Resolution Detector width 20 mm ( 16 x 1.25 mm)• 40 mm (64 X 0.625) or 40 x 0.625 + 12 x 1.25

More important applied mA values

N detectors

CT MULTI-SLICE TECHNOLOGY

• DOSE• Lower dose with multislice CT than with single slice CT.

X-ray beam width < detector width (80 to 90 %)Dose reduction software

• DLP values increase because of larger collimation (40 mm) ; L acquisition > L required

• To compensate for the increase of noise due to the pitch values, the systems increase the mA station ==> constant dose.

effective mA concept

• CTDI is measured in the same conditions than for single slice CT machine.

CT MULTI-SLICE TECHNOLOGY

0

50

100

150

200

250

300

0.0 100.0 200.0 300.0 400.0 500.0

100% 100% 55% 55%

40% 40%

mA

sm

As

mA ConstantsmA ConstantsZ Modulation - Auto mA Z Modulation - Auto mA

XYZ ModulationXYZ Modulation

mA = function of (Image quality needed, tissues attenuation) Optimization of image noise

DOSE MODULATION

zD

fQ

2

3

• Quality of the image– Low noise– Good resolution– Sub-millimeter slices– Low dose

• Image Q factor suggested by « Impact »• f spatial resolution (MTF pl/mm)• σ noise• Z slice thickness (mm)• D dose (CTDI vol)

IMAGE QUALITY FACTOR

(CTDI) and effective dose for different CT examinations (EUR 16262)

Region Head Thorax Abdomen Pelvis

Length of examined Area (mm)

160 320 300 160

Slice thickness (mm) 5

10 5 3

Time (s) 32 32 40 40

Current (A) 210 210 165 165

Organ Eye Lens Lungs Liver Bladder

Organ dose (mSv) 28.1 23.3 12.9 13.3

Effective Dose (mSv) 1,1 6,7 4,3 2,7

PROPOSED REFERENCE DOSE VALUES

mAs VARIATION (SLICE THIKNESS OF 5 mm)

0100200300400500

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17

mAs

French Survey carried out in 2004

0

0,1

0,2

0,3

0,4

0,5

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17

mGy/mAs VARIATION (SLICE THIKNESS OF 5 mm)


EFFECTIVE DOSE COMPARISON (mGy)

0

2

4

6

8

10

12

CHEST ABDOMEN

Axial

helical


EFFECTIVE DOSE (abdomen-pelvis)

mSv

0

5

10

15

20

25

30

helicalaxial

mean

max

meanmin


Routine CT examinations on the basis of absorbed dose to air (EUR 16262 )

Examination Reference dose valueCTDIw (mGy) DLP (mGy cm)

Routine heada 60 1050

Face and sinusesa 35 360

Vertebral traumab 70 460

Routine chestb 30 650

HRCT of lungb 35 280

Routine abdomenb 35 780

Liver and spleenb 35 900

Routine pelvisb 35 570

Osseous pelvisb 25 520

a. Data relate to head phantom (PMMA, 16 cm diameter)

b. Data relate to body phantom (PMMA, 32 cm diameter)

PROPOSED REFERENCE DOSE VALUES

QUALITY CONTROL

Example of QC Test periodicity :

QC Test Acceptance Daily Monthly Annually Mechanic * *

Noise * * Uniformity * *

Low Contrast detectability

* *

Spatial Resolution * * Contrast scale

linearity * *

Slice Thickness * * Dose * *

ct from past to future carlo maccia medical physicist caats 43 bd du maréchal joffre –...

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