ct from past to future carlo maccia medical physicist caats 43 bd du maréchal joffre –...
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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
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
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
IMAGE RESOLUTION (influencing factors)
• Acquisition• Number of projections (floating focal spot
technique)
IMAGE RESOLUTION (influencing factors)
• 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
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
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)
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.
COMPUTED TOMOGRAPHY DOSE INDEX (CTDI)
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
COMPUTED TOMOGRAPHY DOSE INDEX (CTDI)
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
REFERENCE DOSE QUANTITIES
• 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
REFERENCE DOSE QUANTITIES
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
REFERENCE DOSE QUANTITIES
• 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
REFERENCE DOSE QUANTITIES
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)
French Survey carried out in 2004
EFFECTIVE DOSE COMPARISON (mGy)
0
2
4
6
8
10
12
CHEST ABDOMEN
Axial
helical
French Survey carried out in 2004
EFFECTIVE DOSE (abdomen-pelvis)
mSv
0
5
10
15
20
25
30
helicalaxial
mean
max
meanmin
French Survey carried out in 2004
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 * *