XI. National Turkish Medical Physics Congress 14-18 November 2007 - Antalya CT from past to future Carlo Maccia Medical Physicist CAATS 43 Bd du Maréchal Joffre – Bourg-La-Reine – FRANCE

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, x is the path length through the attenuating material, and  is the linear attenuation coefficient 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 n2 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

NOTE 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

Predominant design of current commercially 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 and receive signals from rotating parts. In the slip-ring method, an electrical conductive brush moves along a ring-shaped electrically conductive rail. The use of a slip ring permits 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 CT Detector Array and Collimator X-Ray Tube

A Look Inside a Slip Ring CT Note: how most of the electronics is placed on the rotating gantry X-Ray Tube Detector Array Slip Ring

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

Fast + Poor Image quality Fast + Improved Image quality Single source CT Dual 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.

I0   In I1 = I0 e-1w I2 = I1 e-2w = I0e-w(1 + 2) Recall from previous lecture notes I0   In Pixel Position Output Intensity I1 = I0 e-1w 1 I2 = I1 e-2w = I0e-w(1 + 2) 2 n In = I0e-w(1 + 2 + … + n) therefore Ray Sum Value 1 + 2 + 3 + … + n = 1 ln (I0/ In) w 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 Detector Position a projection is defined as the set of ray sums measured in all detectors during a single x-ray tube pulse typically anywhere between 800 - 1000 projections are collected in one 360° tube rotation to reconstruct a single axial image

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

IMAGE INTERPOLATION (SPIRAL CT) PROBLEM The volume scanned in a single rotation differs between the conventional and helical scanning methods. ANSWER Interpolate 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)

CT Fluoroscopy Real Time Guidance (up to 8 fps) Great Image Quality Low Risk Faster Procedures (up to 66% faster than non fluoroscopic procedures) Approx. 80 kVp, 30 mA

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

CT NUMBER CT Number = 1000 p - w = Hounsfield Unit (HU) w The final result of the CT image reconstruction is an accurate estimate of the x-ray absorption 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 linear attenuation 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

Because p and w are dependent on photon energy (keV), HU values depend on the kVp and filtration. Therefore, HU values generated by a CT scanner are approximate and only valid for the effective kVp used to generate the image.

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 generally increases as kVp decreases but is not affected by mAs or scan time. CT Photon Energy Range (120 or 140 kVp) 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.

LOW CONTRAST RESOLUTION Measurement Technique Catphan 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 numbers calculated from pixel values in a predefined region-of-interest (ROI) using an image of a uniform material (usually water). The selected ROI region should be void of objects 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 NOTE: Noise = 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

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/pixel then 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)

ASIDE (noise)2  1 B • D • H • w3 From the above equation : A 4x increase in dose is required for a 2x decrease in σ. Reduced σ will provide better low contrast detectability An 8x dose increase is required for a 2x decrease in W which maintains constant σ. Decreasing W will increase high contrast detectability. A 2x dose increase is required to decrease H by a factor of 2 and keep σ constant. Decreasing both W and H by 2x would require a dose increase of 16x in order to maintain constant σ. Since B depends exponentially on patient thickness, increasing the thickness of the patient or body part being examined by a factor of n will result in a decrease in transmission from B to Bn. If B = 0.05 and n = 2, a dose increase of 20x would be required to maintain σ, H and W. This is one reason why larger pixels are often used for body scans than for head scans.

IMAGE RESOLUTION Spatial resolution is the ability to discriminate between adjacent objects and is a function 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

Measurement Technique 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 Chest Angiography

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)

RADIATION DOSE Single Slice Profile The dose profile in a CT scanner is not uniform along the patient axis and may vary within any irradiated section. Single Slice Profile Slice Thickness (T) defined at FWHM NOTE: because of scattered x-rays, the CT section dose profile is not perfectly square but has tails that extend beyond the section edges. Tissues beyond the section are thus exposed to radiation

COMPUTED TOMOGRAPHY DOSE INDEX (CTDI) 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).

Measurement principle COMPUTED TOMOGRAPHY DOSE INDEX (CTDI) Measurement principle Mean Dose. Ionization Chamber Aire= e x CTDI CTDI Aire e e e n x e Z mm

COMPUTED TOMOGRAPHY DOSE INDEX (CTDI) 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.

HETEROGENEITY OF DOSE PROFILES Air 1 cm Centre Ideal

) ( COMPUTED TOMOGRAPHY DOSE INDEX (CTDI) 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 ) ( CTDI 3 2 + 1 C = p 100, c w n

C CTDI = × REFERENCE DOSE QUANTITIES 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 = w n ×

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

å C N T CTDI = DLP × 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. C N T CTDI = DLP w n × å i

å In the case of helical (spiral) scanning [mGy • cm] : t A T CTDI = 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. å t A T CTDI = DLP w n i ×

Scan: 15 slices, 10 mm slice thickness and 10 mm slice increment RADIATION DOSE When continuous sections are scanned, the cumulative radiation dose in a section may be as high as twice the radiation dose associated with a single section. Multiple Scan Average Dose (MSAD) Scan: 15 slices, 10 mm slice thickness and 10 mm slice increment

ò (z)dz D = MSAD 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 I N, 2 + - 1 ò Ι

Multiple Scan Average Dose (z)dz D = MSAD I N, 2 + - 1 ò Ι e Z mm T Pitch =1 ; CTDI=MSAD I MSAD : dose delivered while scanning with non adjacent slices (axial mode)

CT MULTI-SLICE TECHNOLOGY N detectors 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

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.

DOSE MODULATION mA = function of (Image quality needed, tissues attenuation) Optimization of image noise 100% 55% 40% mAs mA Constants Z Modulation - Auto mA XYZ Modulation

IMAGE QUALITY FACTOR 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)

PROPOSED REFERENCE DOSE VALUES (CTDI) and effective dose for different CT examinations (EUR 16262)   Region Head Thorax Abdomen Pelvis Length of examined Area (mm) 160 320 300 Slice thickness (mm) 5   10 3 Time (s) 32 40 Current (A) 210 165 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

mAs VARIATION (SLICE THIKNESS OF 5 mm) French Survey carried out in 2004

mGy/mAs VARIATION (SLICE THIKNESS OF 5 mm) French Survey carried out in 2004

EFFECTIVE DOSE COMPARISON (mGy) French Survey carried out in 2004

EFFECTIVE DOSE (abdomen-pelvis) French Survey carried out in 2004

PROPOSED REFERENCE DOSE VALUES Routine CT examinations on the basis of absorbed dose to air (EUR 16262 ) Examination Reference dose value CTDIw (mGy) DLP (mGy cm) Routine heada 60 1050 Face and sinusesa 35 360 Vertebral traumab 70 460 Routine chestb 30 650 HRCT of lungb 280 Routine abdomenb 780 Liver and spleenb 900 Routine pelvisb 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)

QUALITY CONTROL Example of QC Test periodicity :