Physics of Medical Imaging 1 - Ionizing Radiation - PDF

PHYSICS OF MEDICAL IMAGING 1 - IONIZING RADIATION PLANAR IMAGING (RADIOGRAPHY/FLUOROSCOPY) M.Sc. MEDICAL PHYSICS Dr. KHALID IBRAHIM HUSSEIN · · 35 Mc 15 long > - 350-1 mark Per a 30-5marks & answer > - e It Per COMPONENTS OF AN IMAGING SYSTEM Y The Principal components of a system for X ray projec...

PHYSICS OF MEDICAL IMAGING 1 - IONIZING RADIATION PLANAR IMAGING (RADIOGRAPHY/FLUOROSCOPY) M.Sc. MEDICAL PHYSICS Dr. KHALID IBRAHIM HUSSEIN · · 35 Mc 15 long > - 350-1 mark Per a 30-5marks & answer > - e It Per COMPONENTS OF AN IMAGING SYSTEM Y The Principal components of a system for X ray projection radiography: Components such as shaped filtration, compression devices or restraining devices may be added for special cases IAEA Grid and AEC are optional COMPONENTS OF AN IMAGING SYSTEM  For an ideal imaging task with the ideal x-ray spectrum (Monochromatic) : Contrast, C = ∆I/I X where I is the X ray intensity in a background region and ∆I is the difference in X ray intensity for a small detail  The X ray intensity is related to thickness by the attenuation law, therefore the Primary Contrast: where: xd: thickness of the detail µd: linear attenuation coefficient of the detail IAEA µb: linear attenuation coefficient of the background material COMPONENTS OF AN IMAGING SYSTEM X  The relationship between the Linear Attenuation Coefficient and the Mass Attenuation Coefficient tells us that contrast exists for details that differ in:  Mass Attenuation Coefficient and/or Density  Contrast will depend on the Thickness of the detail not the thickness of surrounding tissue  Contrast is Inversely related to the kV setting Since values of µ reduce as photon energy increases  Thus kV may be considered to be the contrast control where contrast is strictly the detail contrast IAEA COMPONENTS OF AN IMAGING SYSTEM Energy absorbed in a small region of the image receptor: Ep: energy absorbed due to primary rays Es: energy absorbed due to secondary rays Scatter may be quantified by: Scatter Fraction SF=Es/(Ep+Es) Scatter to Primary Ratio SP=Es/Ep X The relationship between the two is IAEA SF = ((SP-1)+1)-1 COMPONENTS OF AN IMAGING SYSTEM In the presence of scattered radiation, the Primary Contrast equation becomes X Minimization of scatter is therefore important IAEA · GEOMETRY OF PROJECTION RADIOGRAPHY  The Primary Effect of projection radiography is to record an image of a 3D object (the patient) in 2D, resulting in superposition of the anatomy along each ray - -  This leads to a number of effects that need to be considered in: the Design of equipment the Production of the images and their Interpretation  In particular, for each projection there will be a region of clinical interest, Somewhere between the entrance and exit IAEA of the region to be imaged surface EFFECTS OF PROJECTION GEOMETRY A Geometrical Distortion - Position  All objects are magnified by an * amount related to the OID ⑭ I  The further away from the OID the larger the object appears  In diagram all objects A, B and C Y are the Same size, but they appear progressively larger due to differences in position Effect of depth of objects on their projected size EFFECTS OF PROJECTION GEOMETRY Geometrical Distortion - Shape Tilted object is shown projected at a range of angles, illustrating the increasing degree of foreshortening as the angle increases Effect of angulation on projected length of an angled object EFFECTS OF PROJECTION GEOMETRY Geometrical Unsharpness WIdeal image Sharpness would be produced by a Point Source - m 1XX The spatial resolution in such a case being limited # by the image receptor factors such as ?  Phosphor layer Thickness, · I  Lateral Spread of light in scintillators, and the ~  Image Matrix - & & · 1  The Spatial Resolution depends on the focal spot size. #  Typically the fine focal spots are 0.3-1.0 mm, but must use lower& mAs to protect the tube from heating effects. EFFECTS OF PROJECTION GEOMETRY Geometrical Unsharpness (Ug)  the magnification (m): #I 𝑋 𝑆𝐼𝐷 𝑚= 𝑆𝑂𝐷 where SID is the Source-Image Distance SOD is the Source-Object Distance OID is the Object-Image Distance Ug 𝑂𝐼𝐷 𝑆𝐼𝐷 − 𝑆𝑂𝐷 𝑈 =𝑋. =𝑋. = 𝑋. (𝑚 − 1) 𝑆𝑂𝐷 𝑆𝑂𝐷 EFFECTS OF PROJECTION GEOMETRY N  # ~ Geometrical Unsharpness Optimization of projection radiographs involves choosing an appropriate focal spot size *  This requires a Compromise between the exposure time and the resolution #  For Example, a very small focal spot will provide good - ~ spatial resolution, but only permit a low tube current, therefore requiring a long exposure time, leading to increased risk of um motion blur MAGNIFICATION IMAGING A M  Magnification is achieved by increasing the OID which generally requires an increase in the FID as well. 2  The actual magnification achieved varies with depth in the patient. 1 Example: Patient thickness isO 20 cm, th FID 140 cm and the FSD 80 cm the magnification varies between 1.4 at the Exit side of the patient to 1.75 at the Entrance side. ~ MAGNIFICATION IMAGING Magnification requires employment of a Larger image receptor z For large body regions this may Not be possible *The use of magnification has consequences for:  Dose  Spatial Resolution and  SNR ~ MAGNIFICATION IMAGING Dose - A number of Effects occur when increasing the OID ⑭ There is a substantial reduction in the Scatter Fraction at the image receptor, because the scattered rays are generally directed away from the receptor nAS To maintain the dose to the image receptor, an increase in the / mAs and hence the patient dose would be required * - - m s e i e n m e - MAGNIFICATION IMAGING X Unsharpness wil - &  An increase in the OID leads to a reduction in image sharpness due to the geometric blur of the focal spot #  Use of magnification techniques requires a Significant Reduction in focal spot size compared to contact methods - - MAGNIFICATION IMAGING - Unsharpness Improvement in the overall sharpness of the complete system is generally because of the increase in size of the image compared to the Unsharpness of the image receptor From effects such as:  Light Spread for screen-film systems and the  Pixel Size for digital systems Magnification can therefore Improve Spatial Resolution, compared to the result of a simple zoom of a digital image which enlarges the pixels as well as the image TECHNIQUE SELECTION -  With Screen-Film systems, technique selection is relatively straightforward:  The choice of kV setting is based largely on the required contrast  the mAs is chosen to produce a suitable optical density for the region of clinical interest.  With Digital systems, the direct link between technique setting and image appearance has been lost, making correct technique selection much more difficult TECHNIQUE SELECTION - Effect of Tube Voltage on Contrast, Noise & Dose To determine whether a detail will be detectable in the image, Noise must be considered The primary Source of noise is generally the random arrival of photons at the image receptor, a Poisson process From Rose’s expression, the number of detected photons required per unit area, to image a detail of size d and contrast C with a signal to noise ratio of k, is: N = k2/C2d2 The value of k is often taken to be 5 TECHNIQUE SELECTION - Effect of Tube Voltage on Contrast, Noise & Dose As C is increased, the number of photons required at the image receptor is reduced so that a Reduction in kV will Produce an Image of Satisfactory Quality at a Lower Image Receptor Dose. IAEA - TECHNIQUE SELECTION Technique Selection & 15% Rule However, This Reduction in kV will Require an Increase in the mAs, Leading to an Increase in Patient Dose  The patient dose (Ki), is proportional to mAs and approx to kV2.  The penetration through the patient is proportional to kV3.  Total dose to the image receptor (detector Signal) depends approximately on kV5, and is linear with mAs.  So If the kVp reduce by 15%, the new mAs will be: 𝑚𝐴𝑠 = 𝑚𝐴𝑠 = 𝑚𝐴𝑠 1.15 = 𝑚𝐴𝑠 × 2 TECHNIQUE SELECTION ~ For Example A Lateral Cervical spine radiograph was produced using 32 mAs and 80 kVp at 180cm. The C7-T1 area is not penetrated well and the image needs to be repeated. The kVp is being increased to 92 kVp, what new mAs should be used to maintain the original exposure? 1. The mAs will be increase or decrease? 2. What will be the new value of mAs? 3. Explain the your result? TECHNIQUE SELECTION Y  For all Digital systems the choice of suitable kV and mAs combinations requires that for each projection:  the kV and mAs produce the correct value of the Exposure Index (EI).  the maximum value of the kV is chosen that will allow diagnostically acceptable CNR.  For screen-film imaging this also requires matching the Dynamic Range of the image receptor system to the range of the input signal. NOISE X  X-ray noise is the random variation of the x-ray photons on an x-ray image. That is because of the level of distribution of darker and lighter pixels.  The noise directly contribute to deceasing x-ray image quality. FLUOROSCOPIC EQUIPMENT ~ The Fluoroscopic Imaging Chain  Real time imaging of dynamic things.  Has much longer patient exposure times with lower currents (mAs), typically 1-5 mA.  Build in Grids (10:1 grid ratio). Difer part X-rays Light Input phosphor Electrons Photocathode Light Electron optics & output phosphor FLUOROSCOPIC EQUIPMENT The Fluoroscopic Imaging Chain Input Phosphor converts: X-Rays to Light Most commonly used phosphor: CsI(Tl) crystals grown in a dense needle-like structure - prevents lateral light spread ~ Top portion of a ~750 micron thick film of CsI(Tl) demonstrating well separated columnar growth FLUOROSCOPIC EQUIPMENT The Fluoroscopic Imaging Chain Photocathode converts: Light to Electrons Electrons:  Released through photoelectric effect  Repulsed from photocathode  Accelerated towards anode by 25-30 kV - FLUOROSCOPIC EQUIPMENT The Fluoroscopic Imaging Chain Output Phosphor converts: Electrons to Light Electron beam focused by electrodes onto a thin powder phosphor e.g. ZnCdS:Ag (P20) X FLUOROSCOPIC EQUIPMENT The Fluoroscopic Imaging Chain X Optical System couples XRII to video camera includes:  Collimating Lens to shape the divergent light from the Output Phosphor  Aperture to limit the amount of light reaching the video camera  Lens to focus the image onto the video camera w FLUOROSCOPIC EQUIPMENT The Fluoroscopic Imaging Chain Brightness Gain : gain of output light relative to input light.  Brightness Gain = Minification Gain x Flux Gain Minification Gain is the increase in image brightness due to the decrease in area of the output to input phosphor (~100). Flux Gain: number of light photons emitted from output phosphor relative to input phosphor (~50). Minification Gain = Flux Gain = 𝑫𝒊 𝟐 ( ). 𝑫𝒐 FLUOROSCOPIC EQUIPMENT The Fluoroscopic Imaging Chain X  Video Camera captures the XRII output image, and converts it to an analogue electrical signal.  Older Video Cameras - Photoconductive target scanned by electron beam  Modern Video Cameras - Charge-Coupled Device (CCD) FLUOROSCOPIC EQUIPMENT The Fluoroscopic Imaging Chain & Photoconductive Video Cameras  Resistivity of the photoconductive target changes based on the amount of light striking it creating a Latent Image of the XRII output phosphor.  As the Electron Beam is scanned rapidly across the target, its intensity is modulated by this latent image  The resulting small current is integrated over large resistance and converted to a voltage that is amplified. FLUOROSCOPIC EQUIPMENT X The Fluoroscopic Imaging Chain Photoconductive Video Cameras  Analogue video waveform can be displayed directly on a video monitor  Waveform can also be digitized using an ADC Important ADC characteristics include:  Bit Depth  Sampling Rate FLUOROSCOPIC EQUIPMENT The Fluoroscopic Imaging Chain w CCD Video Cameras  A Solid-State Device composed of many discrete photoconducting cells.  Light from the Output Phosphor is converted to electrons in an Amorphous Silicon photoconducting layer  The electrons are stored in Potential Wells created by applying a voltage between rows and columns of cells IAEA ~ FLUOROSCOPIC EQUIPMENT The Fluoroscopic Imaging Chain CCD Video Cameras  Stored charge that has accumulated during an exposure is read out using parallel and serial Shift Registers that move charge from column to column and row to row.  This creates an analog signal that is amplified and output as a video signal or digitized directly FLUOROSCOPIC EQUIPMENT The Fluoroscopic Imaging Chain Flat Panel Image Receptors Replacing XRIIs in modern systems Advantages include:  Larger FOV Size  Less bulky Profile  Absence of Image Distortions, and a  Higher QDE at moderate to high IAKR Flat panels broaden applications to include Rotational Angiography and Cone-Beam CT. j FLUOROSCOPIC EQUIPMENT The Fluoroscopic Imaging Chain Flat Panel Image Receptors Suffer from Additive noise sources and therefore perform poorly compared to XRIIs at low IAKR Typical IAKR for fluoroscopic imaging with a full-FOV flatpanel receptor (30 cm x 40 cm) range from 27-50 µGy/min ~ ADJUNCT IMAGING MODES S Digital Subtraction Angiography Is type of fluoroscopy technique used in interventional radiology to visualize blood vessels in dense tissue environment. DSA is a technique in which sequential (Fill) images that include a contrast agent are subtracted from a pre-contrast image (Mask) image that includes only the anatomical background This subtraction reduces Anatomical Noise and increases the contrast of the blood vessels in the subtracted images ADJUNCT IMAGING MODES X Digital Subtraction Angiography Major source of Artefacts in DSA: patient motion between the capture of the mask and fill images These motion artefacts can obscure contrast-filled vessels These types of artefacts can be reduced retrospectively in some cases through the use of processing techniques such as manual or automatic Pixel Shifting of the mask image or Remasking through the selection of a different mask frame for subtraction Digital Subtraction Angiography L Image formation  Temporal subtraction and energy subtraction: Number of computer-assisted techniques whereby an image obtain at one time is subtracted from an image obtained at a later time. IAEA Digital Subtraction Angiography X Image formation  Energy subtraction: Uses two different x-ray beams alternately to provide subtraction image that results form differences in photoelectric interaction. IAEA ADJUNCT IMAGING MODES Digital Subtraction Angiography mode ~ Roadmapping: An Adjunct imaging mode used to create a Map of vascular anatomy that aids the Navigation of catheters. Peripheral Runoff Imaging: Follows a Bolus of contrast as it travels from the injection site into the peripheral vasculature, most often in the legs : Rotational Angiography An adjunct imaging mode used most often in vascular, interventional, and neurointerventional radiology. A series of basis images are acquired as a C-arm rotates around the patient X-RAY IMAGE FORMATION Technique Selection ~ Automatic Exposure Control (AEC)  Even with the most skilled of practitioners, manual setting of technique factors results in inconsistent exposures, so that:  Optical densities vary in Screen-Film imaging and  Image noise levels vary with Digital systems  In addition a number of rejects and repeats are unavoidable due to exposure errors  AEC Systems are intended to increase exposure consistency and reduce reject and repeat rates X-RAY IMAGE FORMATION Technique Selection ~ Automatic Exposure Control (AEC) The Principle is to  measure the X ray Flux at the image receptor and to  Terminate the exposure when sufficient Energy has been absorbed Anti Scatter Grids X  The scatter component of the image may be considered to consist of the primary image convolved with a Scatter Spread Function which gives a highly blurred version of the image  The resulting image may be considered to be the Sum of these two images  Efforts are being made to employ this idea for Computerized Scatter Removal, rather than using grids or other methods  This approach is complicated because the scatter spread function is not shift invariant Anti Scatter Grids X In the absence of computerized methods, the use of AntiScatter Grids is routine for the vast majority of radiographic projections, apart from those of the extremities Grids vary greatly in terms of the  Degree of scatter rejection, and in the  Increase in dose to the patient that their use requires All are designed to Allow a large proportion of the primary photons to reach the image receptor whilst Removing a good proportion of the scattered photons from the radiation field Anti Scatter Grids Grid Construction ~ Construction and principle of operation of a focused anti-scatter grid (not to scale) The ratio of the height of the lead strips to the width of the interspace material is known as the Grid Ratio, which is therefore given by: Anti Scatter Grids & Measures of Grid Performance Definitions The Grid Factor or Bucky Factor (BF) is the dose increase factor associated with the use of the grid: The Contrast Improvement Factor is given by Anti Scatter Grids Grid Artefacts & Alignment There are several possible Misalignments that will lead to Artefacts in projection images IAEA A & MAMMOGRAPHY Objective: To familiarize the student with the requirements and principles of imaging of the breast using X ray mammography. mea , -59 ↓IAEA - International Atomic Energy Agency TABLE OF CONTENTS 1. 2. 3. 4. 5. 6. 7. 8. 9. Radiological requirements for mammography X ray equipment Image receptors Display of mammograms Breast tomosynthesis Breast CT Computer-aided diagnosis Stereotactic biopsy systems Radiation dose Bibliography RADIOLOGICAL REQUIREMENTS FOR MAMMOGRAPHY  Breast cancer is a major killer of women.  Mortality can be significantly reduced if disease is detected at an early stage  Mammography is a radiographic (X ray) procedure optimized for examination of the breast  Highly effective means of detecting early-stage breast cancer  Mammography is used both for Investigating symptomatic patients (diagnostic mammography) Screening of asymptomatic women (selected age groups)  Other uses Pre-surgical localisation and guidance of biopsies. - RADIOLOGICAL REQUIREMENTS FOR MAMMOGRAPHY  Breast tissues intrinsically lack subject contrast  Low-energy X ray spectrum required  Emphasises compositional differences of the breast Dependence of the linear X ray attenuation coefficient ( µ) on X ray energy. RADIOLOGICAL REQUIREMENTS FOR MAMMOGRAPHY -  Sufficient spatial resolution Details possibly as fine as 50 µm must be adequately visualised  Adequate contrast in image Low-energy X ray spectra  Broad dynamic range Required due to composition of the breast and age-dependent changes in the breast  Lowest absorbed dose compatible with adequate diagnostic image quality X RAY EQUIPMENT Schematic of a mammography imaging system.  Specialised gantry to accommodate the breast Rotation and vertical movement  Specialised beam geometry Improves visualisation of chest wall edge System geometry for image acquisition showing a) correct alignment and b) missed tissue associated with incorrect alignment. X RAY EQUIPMENT Tubes, filters and spectra  X ray tube Rotating anode Dual focus 0.3/0.1 mm Beryllium exit window (low attenuation)  FID (focus image distance) generally in the range 60 to 65 cm  For screen-film mammography optimum beam energy lies between 18 and 23 keV depending on breast thickness and composition Characteristic X rays from molybdenum  and rhodium are suitable Higher energies may be more optimal for digital mammography The geometry of an X ray tube. The perpendicular line abuts the chest wall. The reference axis on a particular system will be specified by the manufacturer. X RAY EQUIPMENT Tubes, filters and spectra  Metallic filters used in mammography  Molybdenum (Mo) filter (30 to 35 µm thick) commonly employed with Mo anode  Filter acts as energy window Greater attenuation of X rays at low energies and at energies above the K-absorption edge of Mo at 20 keV Mo characteristic X rays from the target and X rays of similar energy produced by bremsstrahlung pass through the filter Resultant spectrum enriched with X rays in the range 17 to 20 keV  Higher energies are desirable for imaging thick, dense breasts Use of Mo/Rh (molybdenum/rhodium) and Rh/Rh target/filter combinations X RAY EQUIPMENT Tubes, filters and spectra Examples of mammographic X ray spectra. X RAY EQUIPMENT Compression & Grids  Reduces superposition of tissues  Decreases ratio of scattered to transmitted radiation reaching the image receptor  Compressed breast provides lower overall attenuation allowing radiation dose to be reduced  Compressed breast provides more uniform attenuation over the image reducing the exposure range which must be recorded  Scattered radiation reduces image quality  Use of grid significantly decreases ratio of scattered to transmitted radiation reaching the image receptor  Bucky factor (increase in dose due to use of grid) can be as large as 2 to 3 Justified by improvement in image quality X RAY EQUIPMENT Automatic exposure control  All modern mammography units are equipped with automatic exposure control (AEC)  Essential in order to provide the optimum dose to the image receptor Target optical density for screen-film mammography Target SNR (signal-to-noise ratio) or preferably SDNR (signaldifference-to- noise ratio) for digital mammography  For digital mammography the digital detector can act as the AEC sensor X RAY EQUIPMENT Magnification mammography  Magnification mammography can be used to improve the diagnostic quality of the image  Breast supported above the image receptor Focus object distance reduced Object to image receptor distance increased Magnification results  Benefits of magnification mammography Increased SNR Improved spatial resolution Dose-efficient scatter rejection. X RAY EQUIPMENT Magnification mammography  Spatial resolution in magnification mammography is limited by focal spot size Use of a small spot (typically 0.1 mm) is critical  As the breast is closer to the X ray source in magnification mammography Dose to breast increases (compared to contact mammography) Air gap between the breast and image receptor provides some scatter rejection Anti-scatter grids not employed for magnification (partially offsets increase in dose) IMAGE RECEPTORS  Screen-film mammography: Single back intensifying screens used with single emulsion radiographic film enclosed in a light-proof cassette. High resolution fluorescent intensifying screen  Digital mammography Area detectors Indirect detectors Direct detectors Photo-stimulable phosphors (computed radiography or CR) Scanning detectors Photon counting detectors IMAGE RECEPTORS Screen-film mammography By minimizing the distance that the light must travel before being collected by the film, blurring due to lateral spreading is reduced Spatial resolution is improved Typical phosphor used for screen-film mammography is (Gd2O2S:Tb). gadolinium oxysulphide Photographic film emulsion for mammography is matched to be sensitive to: Spectrum of light emitted from the particular phosphor screen Range of X ray fluence exiting the breast Configuration for a mammographic screen-film image receptor. A single-emulsion radiographic film is held in close contact with a fluorescent screen in a light-proof cassette. IMAGE RECEPTORS Digital mammography  Digital mammography introduced commercially in 2000 Able to overcome many of the technical limitations of screen-film mammography  In digital mammography, image acquisition, processing, display, and storage are performed independently, allowing optimisation of each  Acquisition performed with low-noise X ray detectors with wide dynamic range  As the image is stored digitally: It can be displayed with contrast independent of the detector properties Image processing techniques that are found to be useful can be applied prior to image display IMAGE RECEPTORS Digital mammography  Challenges in creating a digital mammography system with improved performance are mainly related to the X ray detector and the display device.  The detector should have the following characteristics: Efficient absorption of the incident radiation beam Linear or logarithmic response over a wide range of incident radiation intensity Low intrinsic noise to ensure that images are X ray quantum noise limited Limiting spatial resolution of the order of 5 to10 cycles/mm Can provide at least an 18x24 cm and preferably a 24x30 cm field size IMAGE RECEPTORS Digital mammography  Two main approaches in detector development  Area detectors: based of an amorphous silicon thin-film transistor (TFT) panel Entire image is acquired simultaneously and Fast image acquisition Can be used with conventional design of mammography X ray unit equipped with a grid to reduce scatter  Scanning detectors Image is obtained by scanning the X ray beam and detector across the breast Mechanically complex and longer acquisition times Use relatively simple detector(s) Good intrinsic scatter rejection IMAGE RECEPTORS Digital mammography  This readout system allows the signals from all of the dels to be read in a fraction of a second Fast image display  The needle-like phosphor crystals of CsI behave somewhat like fibre-optics Conduct the light to the photodiodes with less lateral spread than would occur with granular phosphors Allows the thickness of the phosphor to be increased to improve the quantum detection efficiency of the detector without excessive loss of spatial resolution IMAGE RECEPTORS Digital mammography  Computed radiography (CR) systems can be used for mammography  Employs a photo-stimulable phosphor (PSP) plate housed in a light-proof cassette When exposed to X rays, electrons in the crystalline material are excited and subsequently captured by traps in the phosphor After exposure the plate is placed in a reader device and scanned with a laser beam The energy of the laser light stimulates the traps to release the electrons The transition of these electrons through energy levels in the phosphor crystal results in the emission of light The light is collected by a photomultiplier tube, the signal digitised and attributed to a particular pixel in the image IMAGE RECEPTORS Digital mammography  Mammography CR systems differ from the general radiography CR systems in several key areas Mammography CR system is designed for higher spatial resolution Uses a thinner phosphor material and is scanned with finer sampling pitch (typically 50 µm)  However, the result is less signal per pixel  To overcome this limitation various innovations have been developed to improve light coupling and reduce readout noise, including: Dual-sided readout of the phosphor plates Needle-like phosphors which permit the use of thicker detectors having superior quantum detection efficiency IMAGE RECEPTORS Digital mammography  Detector systems discussed so far acquire the image by integrating the signal from a number of X ray quanta absorbed in the detector and digitizing this signal  The image noise from these systems depends on: Poisson X ray quantum fluctuations associated with X ray absorption Additional noise sources associated with the production of the converted electronic signal These noise sources can arise from: Fluctuation in the amount of light produced in a phosphor in response to absorption of an X ray of a particular energy or from the X ray spectrum itself  As an alternative it is possible to count the number of interacting quanta directly, thereby avoiding these additional noise sources IMAGE RECEPTORS Digital mammography  Typically quantum-counting detectors employ a geometry in which the X ray beam is collimated into a slot or multi-slit format and scanned across the breast to acquire the image  The detector can be based on: Solid-state approach (electron-hole pairs are produced in a material such as crystalline silicon) Pressurized gas (the signal is in the form of ions formed in the gas) Collection of charge signal plus amplification produces a pulse for each interacting X ray quantum and pulses are counted to create the signal  As the beam is collimated to irradiate only part of the breast at a time the system has: Good intrinsic scatter rejection without the need for a grid Increased dose efficiency.. DISPLAY OF MAMMOGRAMS  Film mammograms ~ Specially designed transillumination devices: The luminance levels must be appropriate for reading mammograms and sufficient to illuminate areas of interest  Digital mammograms Computer displays and workstations: Display system plays a major role in influencing overall performance of digital mammography Image quality presented to the film reader Ease of image interpretation 2 BREAST TOMOSYNTHESIS  In projection radiography tissue superposition can result in a masking effect Qu  Breast tomosynthesis can provide reconstructed planar images of sections of the breast Can aid in reducing the masking effect Q  Breast tomosynthesis generally based on modified digital mammography systems Planar digital imaging and tomosynthesis ~ Tomosynthesis only ~ 3 BREAST TOMOSYNTHESIS -55  X ray tube pivots about a point  Breast platform remains stationary m  Detector usually stationary but may also move  Series of low-dose projection images (typically 9 to 25) & - ~ & & & - - acquired over a limited range of angles (±7° to ±30°)  X ray spectrum Higher energy typically employed (e.g.W/Al)  X ray tube movement Continuous exposure or series of discrete exposures (“step and shoot”)  Total acquisition time must be minimized Possible image degradation due to patient motion Y BREAST TOMOSYNTHESIS R  Planar cross-sectional images are reconstructed from the projections using filtered back projection or an iterative reconstruction algorithm  Because of the limited range of acquisition angles Projection data do not form a complete set Reconstructed image is not a true 3D representation of the breast Possibility of artefacts in the images 1- Question: Explain the concept of "tissue superposition" and how breast tomosynthesis helps to mitigate this issue.Answer: Tissue superposition refers to the overlapping of different breast tissues in a single 2D image, which can obscure lesions. Breast tomosynthesis creates a series of thin-sliced images, allowing for better differentiation of overlapping structures and improved lesion visibility. 2-Question: Why is it important to minimize total acquisition time in breast tomosynthesis, and how is this achieved?Answer: Minimizing acquisition time is crucial to reduce motion artifacts caused by patient movement. This is achieved by using a higher energy X-ray spectrum (W/Al) for better penetration and faster imaging. 3-Question: What are the differences between filtered back projection and iterative reconstruction algorithms used in breast tomosynthesis? Answer: Filtered back projection is a simpler and faster algorithm but can lead to streak artifacts. Iterative reconstruction is more complex and time-consuming but can provide better image quality with reduced artifacts. 4-Question: What are some of the technical challenges and limitations associated with the reconstruction process in breast tomosynthesis? Answer: The limited range of acquisition angles in breast tomosynthesis results in incomplete projection data, leading to challenges in accurately reconstructing a true 3D representation of the breast. This can result in artifacts in the reconstructed images. 5-Question: Can breast tomosynthesis systems also perform traditional 2D mammography, and if so, which type of system offers this capability? Answer: Yes, planar digital imaging and tomosynthesis systems can perform both traditional 2D mammography and tomosynthesis, while tomosynthesis-only systems are dedicated solely to tomosynthesis. ⑭ Question: Explain how breast tomosynthesis works and its advantages over traditional mammography. Answer: Breast tomosynthesis is a modified form of digital mammography that acquires a series of low-dose projection images at different angles around the breast. These images are then reconstructed into planar cross-sectional images, reducing the tissue superposition (masking effect) seen in traditional mammography. This can lead to improved detection and characterization of breast lesions. Question: Describe the image acquisition process in breast tomosynthesis, including the movement of the X-ray tube and detector, and the typical number of projection images acquired. Answer: In breast tomosynthesis, the X-ray tube pivots around a point while the breast platform remains stationary. The detector is usually stationary but may also move. A series of 9 to 25 low-dose projection images are acquired over a limited range of angles (typically ±7° to ±30°). Question: Discuss the reconstruction process in breast tomosynthesis, including the algorithms used and the limitations due to the limited range of acquisition angles. Answer: Planar cross-sectional images are reconstructed from the projections using filtered back projection or an iterative reconstruction algorithm. Due to the limited range of acquisition angles, the projection data is incomplete, and the reconstructed image is not a true 3D representation of the breast. This can lead to artifacts in the images. Question: Explain why a higher X-ray energy spectrum is typically employed in breast tomosynthesis compared to traditional mammography. Answer: A higher energy X-ray spectrum (e.g., W/Al) is typically employed in breast tomosynthesis to better penetrate the breast tissue and reduce image noise. This is important because the total acquisition time must be minimized to avoid image degradation due to patient motion. Question: What are the two main types of breast tomosynthesis systems, and how do they differ in terms of X-ray tube movement? Answer: The two main types of breast tomosynthesis systems are: Planar digital imaging and tomosynthesis systems: These systems can perform both traditional 2D mammography and tomosynthesis. Tomosynthesis-only systems: These systems are dedicated to tomosynthesis and cannot perform traditional 2D mammography. The X-ray tube movement in breast tomosynthesis can be either continuous exposure or a series of discrete exposures ("step and shoot"). * 1 How can tissue superposition lead to a masking effect in projection radiography? 2 What role does Breast Tomosynthesis play in reducing the masking effect in imaging? 3 Explain the main difference between Breast Tomosynthesis and traditional digital mammography systems. 4 What are the key components that remain stationary during a Breast Tomosynthesis procedure? 5 Describe the process of acquiring projection images in Breast Tomosynthesis in terms of angles and X-ray dose. 6 Why is higher energy typically employed in the X-ray spectrum during Breast Tomosynthesis? 7 Compare and contrast continuous exposure and "step and shoot" methods in X-ray tube movement during imaging. 8 Why is it crucial to minimize the total acquisition time during Breast Tomosynthesis imaging? 9 How are planar cross-sectional images reconstructed from projections in Breast Tomosynthesis? 10 What are the limitations of reconstructed images in Breast Tomosynthesis due to the limited range of acquisition angles? 1 Tissue superposition can lead to a masking effect in projection radiography when different structures overlap, making it difficult to distinguish individual details. 2 Breast Tomosynthesis can reduce the masking effect in imaging by providing reconstructed planar images of sections of the breast, which helps separate overlapping structures. 3 The main difference between Breast Tomosynthesis and traditional digital mammography systems is that Tomosynthesis provides both planar digital imaging and reconstructed tomographic images, while traditional systems only offer planar imaging. 4 The key components that remain stationary during a Breast Tomosynthesis procedure are the breast platform and the detector (which may also move). 5 In Breast Tomosynthesis, a series of low-dose projection images are acquired over a limited range of angles (typically ±7° to ±30°) to create a 3D image. 6 Higher energy is typically employed in the X-ray spectrum during Breast Tomosynthesis to penetrate the breast tissue and reduce scatter, improving image quality. 7 Continuous exposure involves the X-ray tube moving continuously during imaging, while "step and shoot" method has the tube move in discrete exposures, reducing radiation dose. 8 It is crucial to minimize the total acquisition time during Breast Tomosynthesis imaging to avoid image degradation due to patient motion. 9 Planar cross-sectional images are reconstructed from projections in Breast Tomosynthesis using filtered back projection or an iterative reconstruction algorithm. 10 The limitations of reconstructed images in Breast Tomosynthesis due to the limited range of acquisition angles include artefacts and the reconstructed image not being a true 3D representation of the breast. E BREAST CT  Dedicated breast CT systems have been developed using cone-beam geometry and a flat-panel X- ray detector Data for all of the CT slices acquired simultaneously Rapid image acquisition Pixel dimensions substantially larger than for digital mammography or tomosynthesis Large number of projections To keep doses at an acceptable level images are generally acquired at a higher tube voltage (50 to 80 kV) Very low dose per projection can result in noisy images  Current designs provide a dedicated prone imaging table  Breast CT can be performed without the need to compress the breast. Y COMPUTER-AIDED DIAGNOSIS  Computer-aided diagnosis or computer-aided detection (CAD) systems are designed to assist the film reader in detecting breast cancers  Computer system with sophisticated pattern recognition software Natural adjunct to digital mammography Screen-film mammograms must be digitised (scanned)  Identifies “suspicious” features and alerts image reader Does not replace the image reader  CAD algorithms must be trained using sets of mammograms for which the presence or absence of cancers is known 7 COMPUTER-AIDED DIAGNOSIS  Results of CAD are conveyed to the film reader by means of an image annotated to show the computer detections Different symbols for different lesions  Main use in screening mammography # Double reading has been shown to increase the cancer detection rate CAD has the potential to be a cost-effective alternative to double reading CAD has the potential to: Reduce the number of missed cancers Reduce the variability between film readers & STEREOTACTIC BIOPSY SYSTEMS  Stereotactic procedures are used to investigate suspicious mammographic or clinical findings without the need for surgical biopsies Reduced patient risk, discomfort and cost  In stereotactic biopsies, the gantry of a mammography unit has the facility to allow a pair of angulated views of the breast (typically at ± 15° from normal incidence) to be obtained  From measurements obtained from these images, the three-dimensional location of a suspicious lesion is determined and a needle equipped with a spring-loaded cutting device can be accurately placed in the breast to obtain tissue samples a STEREOTACTIC BIOPSY SYSTEMS These systems may use small-format (e.g. 5 cm x 5 cm) digital detectors or fullfield digital detectors The geometry for stereotactic breast biopsy is shown. The X ray tube is rotated about the breast to produce two views. The Z-depth of an object can be determined by the lateral (X) displacement observed between the two views. RADIATION DOSE so a  Three dosimetric quantities used in mammography Incident air kerma (IAK), Ki Entrance surface air kerma Ke Mean glandular dose (MGD, mean dose to the glandular tissue of the breast), DG  MGD is the primary quantity of interest related to the risk of radiation induced cancer in breast imaging  MGD is calculated using factors obtained experimentally or by Monte Carlo radiation transport calculations  IAK (measured) converted to MGD for a breast of specific composition and size Conversion coefficients are tabulated in various publications (including IAEA Technical Report TRS-457) RADIATION DOSE sur X  Dose in mammography depends on the size and composition of the breast as well as the imaging device and exposure settings selected.  K increases as the thickness and/or density of the breast e increase resulting in a corresponding increase in MGD.  Increase in beam energy (tube voltage, target/material combination) will decreases the radiation dose BUT Image contrast will be reduced and at some point this will become unacceptable RADIATION DOSE u S &  In digital mammography goal is to maintain a target SDNR     at the detector. K increases as the thickness and/or density of the breast increase resulting in a corresponding increase in MGD. Increase in beam energy (tube voltage, target/material combination) will decreases the radiation dose. On a digital system where contrast can be adjusted during image display, an acceptable compromise can be achieved at a higher energy than with screen-film imaging. This allows the advantage of a greater relative decrease of dose compared to film for large and/or dense breasts e BIBLIOGRAPHY  AMERICAN COLLEGE OF RADIOLOGY, Stereotactic breast biopsy quality control manual, American College of Radiology, Reston, VA (1999).  AMERICAN COLLEGE OF RADIOLOGY, Mammography quality control manual, American College of Radiology, Reston, VA (1999).  BICK, U., DIEKMANN, F., Digital mammography, Springer, Heidelberg, Germany (2010).  EUROPEAN COMMISSION, European Guideline for Quality Assurance in Mammography Screening, Office for Official Publications of the European Communities Rep. V4.0 Luxembourg (2006). http://www.euref.org. BIBLIOGRAPHY  PRESTON, D.L., et al., Radiation effects on breast cancer risk: a pooled analysis of eight cohorts, Radiat Res 158 2 (2002) 220-35. http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?cmd=Retriev e&db=PubMed&dopt=Citation&list_uids=12105993.  FERLAY, J., et al., "Cancer Incidence and Mortality Worldwide in 2008: IARC CancerBase No. 10", GLOBOCAN 2008 (Proc. Conf. Lyon, France, 2008), International Agency for Research on Cancer, World Health Organization, http://globocan.iarc.fr. BIBLIOGRAPHY  INTERNATIONAL ATOMIC ENERGY AGENCY, Quality Assurance Programme for Screen-Film Mammography, Human Health Series, 2, IAEA Vienna (2009). http://wwwpub.iaea.org/MTCD/publications/PDF/Pub1381_ web.pdf.  INTERNATIONAL ATOMIC ENERGY AGENCY, Quality Assurance Programme for Digital Mammography, Human Health Series 17, IAEA Vienna (2011).  INTERNATIONAL COMMISSION ON RADIATION UNITS AND MEASUREMENTS, Mammography - Assessment of image quality, ICRU Rep. 82, Journal of the ICRU Vol. 9 No. 2 Bethesda, MD (2009). Computed Tomography TABLE OF CONTENTS 1. 2. 3. 4. Introduction CT principles The CT imaging system Image reconstruction and processing INTRODUCTION  Clinical Computed Tomography (CT) was introduced in X 1971 - limited to axial imaging of the brain in neuroradiology  Nowadays dedicated CT scanners are available for special clinical applications, such as For radiotherapy planning - these CT scanners offer an extra wide bore, allowing the CT scans to be made with a large field of view. The integration of CT scanners in multi modality imaging applications, for example by integration of a CT scanner with a PET scanner or a SPECT scanner. CT PRINCIPLES X X-ray projection, attenuation and acquisition of transmission profiles  The purpose of a computed tomography acquisition is to measure x ray transmission through a patient for a large number of views. CT PRINCIPLES X-ray projection, attenuation and acquisition of transmission profiles X  As an X ray beam is transmitted through the patient, different tissues are encountered with different linear attenuation coefficients.  The intensity of the attenuated X ray beam, transmitted a distance d, can be expressed as: The linear attenuation coefficient (µ) depends on the composition of the material, the density of the material, and the photon energy CT PRINCIPLES Hounsfield Units Meas -  In CT the matrix of reconstructed linear attenuation coefficients (µmaterial) is transformed into a corresponding matrix of Hounsfield units (HUmaterial), where the HU scale is expressed relative to the linear attenuation coefficient of water at room temperature (µwater): HUmaterial  µmaterial µwater 1000 µwater  It can be seen that HUwater = 0 as (µmaterial = µwater), HUair = -1000 as (µmaterial = 0) HU=1 is associated with 0.1% of the linear attenuation coefficient of water. CT NUMBER WINDOW /  Hounsfield units are usually visualized in an eight bit grey scale offering only 128 grey values.  The display is defined using 4000+ HU 4000+ HU Window level (WL) as CT number of mid-grey Window width (WW) as the number of HU from black -> white WL WW 0 HU 0 HU WL -1000 HU -1000 HU WW CT PRINCIPLES Hounsfield Units  Same image data at different WL and WW 4000+ HU 4000+ HU 0 HU WL WL WW WW -1000 HU -1000 HU WL -593, WW 529 WL -12, WW 400 GSTT Nuclear Medicine 06 THEHistorical CT IMAGING SYSTEM and current acquisition configurations Technological advances, 1985 - 2007 Y X THE CT IMAGING SYSTEM Gantry and table  Electrical power is generally supplied to the rotating gantry through contacts (brushes) from stationary slip rings.  Projection profiles are transmitted from the gantry to a computer usually by wireless communication (or slip ring contacts). IAEA power THE CT IMAGING SYSTEM X Gantry and table  The position of the patient on the table can be head first or feet first supine or prone  The position is usually recorded with the scan data. impactscan.org THE CT IMAGING SYSTEM X Collimation and filtration  The X-ray beam is often referred to as a fan beam where the beam width along the longitudinal axis is small  For multi-slice scanners where the longitudinal beam width is no longer small the X-ray beam is often referred to as ‘cone beam’ IAEA - THE CT IMAGING SYSTEM Collimation and filtration  Schematic figure showing the fan beam, flat and beam shaping (‘bow-tie’) filters  The purpose of the beam shaping filter is to reduce the dynamic range of the signal recorded by the CT detector Reduce the dose to the periphery of the patient Attempt to normalise the beam hardening of the beam – to aid with calibration THE CT IMAGING SYSTEM X Detectors  CT detectors are curved in the axial plane (x-y plane), and rectangular along the longitudinal axis (z-axis).  Essential physical characteristics of the CT detectors are Good detection efficiency Fast response (and little afterglow) Good transparency for the generated light Detector Type Efficiency Xenon gas filled 70% Solid state Approaching 100% X THE CT IMAGING SYSTEM Detectors  Resolution in the reconstructed images depends on The size of detector elements - along the detector arc and the z-axis The angular separation of the projection ImPACT x-y plane z-axis x-y plane THE CT IMAGING SYSTEM v Was Detectors  Detector sizes are the effective size at the iso-centre The minimum number of detector elements should be approximately (2 FOV)/d to achieve a spatial resolution of d in the reconstructed image → ~ 800 detector elements are required to achieve a spatial resolution of 1 mm within a reconstructed image at a field of view of 400 mm x-y plane d FOV  Spatial resolution can be improved by use of the quarter detector shift  It can also be improved by the use of a dynamic focal spot Schematic view of detector sizes ImPACT THE CT IMAGING SYSTEM Detectors  Quarter detector shift By shifting the detector elements by a distance of a quarter of the size of the detector elements, the theoretical achievable spatial resolution becomes twice as good. It is generally implemented in detectors of all CT scanners. no ¼ shift with ¼ shift ImPACT Schematic view of quarter detector shift – X-Y plane THE CT IMAGING SYSTEM Detectors  Dynamic or flying focal spot Focal spot position on anode is rapidly oscillated during gantry rotation, doubling the number of projections. A spatial resolution of ~ 0.6 – 0.9 mm in the axial plane can be achieved Schematic view of dynamic focal spot – X-Y plane THE CT IMAGING SYSTEM Detectors  A multi-slice scanner may be defined by the number of ‘data slices’ it acquires – or by the number of detector rows e.g.GE LightSpeed, four slice scanner has 16x 1.25 mm detectors it can acquire 4 x 1.25 mm, 4 x 2.5 mm, 4 x 3.75 mm, or 4 x 5 mm slices THE CT IMAGING SYSTEM Detectors  Multi-slice or multi-row scanners enabled Thinner slices Longer scan volumes Faster scan volumes  A typical acquisition with a single detector row scanner covered 5 mm.  CT scanners with 4 active detector rows achieved a substantial improvement of the longitudinal resolution. For example, by using 4 active detector rows in a 4 x 1 mm acquisition configuration, the longitudinal spatial resolution improved from 5 mm to 1 mm THE CT IMAGING SYSTEM Detectors  The CT scanners with 16 or 64 active detector rows allowed for acquisitions in for example 16 x 0.5 = 8 mm and 64 x 0.5 = 32 mm configurations. These scanners provided excellent longitudinal spatial resolution, high quality 3D reconstructions, and at the same time reduced scan times.  The CT scanners with up to 64 active detector rows generally cover a scan volume with a helical acquisition with multiple rotations.  The 320 detector row CT scanner covers 160 mm on one rotation, for organs such as the brain or the heart within one rotation. mca IMAGE RECONSTRUCTION AND PROCESSING % :-G) ggs ef General concepts  Techniques for reconstruction include Ei Fast more Simple backprojection > more noise but fast noise Algebraic reconstruction but low slow > Iterative reconstruction is > & gissj Filtered back projection but noise - - jojosa & In summary, while FBP is fast and traditionally used, IR offers improved image quality and lower radiation doses at the cost of increased computational demands. The choice between the two methods may depend on the specific requirements of the imaging task at hand2. 11.4 IMAGE RECONSTRUCTION AND 11.4.1 General concepts PROCESSING X  During a CT scan, numerous measurements of the transmission of X-rays through a patient are acquired at many angles  This is the basis for reconstruction of the CT image. x-ray tube I0 attenuation I(d) detector element x-y plane IAEA Diagnostic ImPACT RadiologyPhysics: a Handbook for Teachers and IMAGE RECONSTRUCTION AND PROCESSING General concepts X  The logarithm of the (inverse) measured normalized transmission, ln(I0/I(d)), yields a linear relationship with the products of µi∆x. IMAGE RECONSTRUCTION AND PROCESSING General concepts X  The figure below shows (a) the X-ray projection under a certain angle (b) leading to one transmission profile  The backprojection distributes the measured signal evenly over the area (c) under the same angle as the projection IMAGE RECONSTRUCTION AND PROCESSING General concepts  Transmission profiles are taken from a large number of angles and backprojected (d) yielding a strongly blurred image A  Mathematics shows that simple backprojection is not sufficient for accurate image reconstruction in CT.  Instead a filtered backprojection must be used It is the standard for image reconstruction in CT. IMAGE RECONSTRUCTION AND PROCESSING General concepts X  Other reconstruction techniques are algebraic or iterative reconstructions.  Algebraic reconstruction solves a number of simultaneous equations.  For example Projections in two horizontal, two vertical, and two diagonal directions yield six projection values. These values can be used to solve an overcomplete set of six equations. The equations can be solved and they yield the 2 x 2 image matrix. IMAGE RECONSTRUCTION AND PROCESSING X General concepts  Algebraic reconstruction in clinical practice is not feasible, due to the large (512 x 512) matrices that are used in medical imaging due to inconsistencies in the equations due to measurement errors and noise. IMAGE RECONSTRUCTION AND PROCESSING Y General concepts  Iterative (statistical) reconstructions are sometimes used These are routinely used in nuclear medicine. They are becoming available for commercial CT scanners  Potential benefits of iterative reconstructions the removal of streak artefacts (particularly when fewer projection angles are used) better performance in low-dose CT acquisitions  Filtered backprojection is the most frequently applied technique for CT reconstruction IMAGE RECONSTRUCTION AND PROCESSING X Object, image and radon space  The three domains associated with the technique of filtered backprojection are a) Object space (linear attenuation values), b) Radon space (projection values recorded under many angles) this domain is also referred to as sinogram space where Cartesian coordinates are used), c) Fourier space which can be derived from object space by a 2D Fourier transform. Object space Radon space Fourier space IMAGE RECONSTRUCTION AND PROCESSING X Object, image and radon space  The figure below illustrates the interrelations between the three domains for one projection angle. (a) One specific projection angle in object space (b) The projection that is recorded by the CT scanner (c) This projection corresponds with one line in Radon space (d)One angulated line in Fourier space is created from a 1-D transform of the recorded line in the sinogram IMAGE RECONSTRUCTION AND PROCESSING Object, image and radon space  The interrelationships between the three domains object space, Radon space, and Fourier space X IMAGE RECONSTRUCTION AND PROCESSING Filtered back projection and other reconstructions X  The mathematical operations that are required for a filtered backprojection consist of four steps. 1. A Fourier transform of Radon space should be performed (requiring many 1D Fourier transforms). 2. Then a high-pass filter should be applied to each one of the 1D Fourier transforms. IMAGE RECONSTRUCTION AND PROCESSING X  The mathematical operations that are required for a filtered backprojection consist of four steps. 3. Next an inverse Fourier transform should be applied to the high pass filtered Fourier transforms in order to obtain a Radon space with modified projection profiles. 4. Subsequently, backprojeOcbtjeioctnspaocefthe filtered profiles yields the reconstruction of the measured object. 1-D FT Object space Radon space Fourier space Radon space Back P. 4. Back projection of filtered profiles 3Fo.urie1r-sDpacFeT High Pass filter IMAGE RECONSTRUCTION AND PROCESSING  Successive filtered backprojections with 1, 2, 4, 8,16, 32, 64, 256, and 1024 Backprojections  This shows how successive filtered backprojections under different angles can be used to achieve a good reconstruction of the space IAEAdomain X IMAGE RECONSTRUCTION AND PROCESSING X  Image space is generally represented on a regular grid The 2D image space is defined as ƒ(x,y), where (x,y) are the Cartesian coordinates in image space.  One 1D projection of the 2D image space with equidistant and parallel rays yields one line in Radon space expressed as the projection p(t,θ), where t is the distance from the projected x-ray to the iso-center and θ is the projection angle IMAGE RECONSTRUCTION AND PROCESSING X  The mathematics of the image reconstruction process, can be expressed compactly in the following equation. |ρ| is the Fourier form of the filter Computed Tomography Part 2 Dr. Khalid Ibrahim Hussein Department of Radiological sciences King Khalid University DR. KHALID IBRAHIM HUSSEIN - KKU TABLE OF CONTENTS 1. 2. 3. CT Acquisition Computed tomography image quality CT Dose DR. KHALID IBRAHIM HUSSEIN - KKU CT ACQUISITION  The actual CT scan is generally preceded by a 2D scan projection radiograph (SPR) to assist in planning the scan  The SPR is acquired with a stationary (non-rotating) X-ray tube, a narrowly collimated fan beam and a moving table.  The X-ray tube is fixed, generally in a position that yields either a frontal or lateral SPR of the patient. DR. KHALID IBRAHIM HUSSEIN - KKU CT ACQUISITION Axial CT scan  An axial CT scan involves an acquisition of transmission profiles with a rotating X-ray tube and a static table.  The complete examination is achieved by translation of the table (“step”) after each axial acquisition (“shoot”) this is referred to as a step and shoot acquisition, sequential or axial scanning Usually, the table translation is equal to the slice thickness so that subsequent axial acquisitions can be reconstructed as contiguous axial images. DR. KHALID IBRAHIM HUSSEIN - KKU ACQUISITION e Helical CT scan  Helical scanning allows for the acquisition of a large volume of interest within one breath hold.  The table translation is generally expressed relative to the (nominal) beam width (in single slice CT this equals the slice width): the ratio of table translation per 360° tube rotation relative to the nominal beam width is in helical CT referred to as the pitch factor. DR. KHALID IBRAHIM HUSSEIN - KKU ACQUISITION MDCT scan  The introduction of fast rotating multislice CT scanners occurred in the late 1990’s – ten years after the introduction of helical CT. This provided the preconditions for new clinical applications  In single slice CT scanners only one linear array of detectors was used The rotation time of single slice CT scanners was 1-2 s, the slice thickness (and nominal beam width) in most clinical applications 5-10 mm.  In multislice scanners 4,16 and 64 adjacent active arrays of detectors were used, enabling the simultaneous measurement of a corresponding large number of transmission profiles. This allows rotation time dropped to well below 1 s, to 0.3-0.4 s. DR. KHALID IBRAHIM HUSSEIN - KKU CT ACQUISITION Cardiac CT  Cardiac scanning requires the cardiac motion to be minimised. Therefore to “freeze” the motion Image during phase of least cardiac motion (generally diastole, or end systole) IAEA DR. KHALID IBRAHIM HUSSEIN - KKU CT ACQUISITION MODES  Retrospective ECG-gated reconstructions A helical scan is performed with an overlapping pitch The cardiac phase selection data is selected retrospectively based on registration of the raw data and the ECG during one or more entire cardiac cycles. To reduce the radiation dose in the phases that are not of interest ECG dose modulation is used. DR. KHALID IBRAHIM HUSSEIN - KKU CT ACQUISITION MODES  Prospective ECG-triggered reconstructions are “step-andshoot” (i.e “axial”) acquisitions. An advantage of such acquisitions is the reduction of patient dose. IAEA DR. KHALID IBRAHIM HUSSEIN - KKU CT ACQUISITION MODES  Retrospectively ECG-triggered Image reconstruction data selected retrospectively. The scan overlapping pitch around 0.2mm.  Full helical scan-data collected at all phases  Flexibility to select data for best phase for cardiac CTA  Use multiple phases for functional studies. IAEA DR. KHALID IBRAHIM HUSSEIN - KKU DR. KHALID IBRAHIM HUSSEIN - KKU CT ACQUISITION MODES Scanning mode Cardiac mode Features “step and shoot” / Axial / Sequence Prospective triggering Padding – to provide greater flexibility of reconstruction Helical Retrospective gating Helical Prospective triggering ECG modulation – to provide reduced dose from constant irradiation. Some margin left for flexibility of reconstruction. Tube current(mA) decreases to a prescribed minimum value outside phase region of interest. Modulation dose = 0 11.5 ACQUISITION 11.5.6 CT fluoroscopy and interventional procedures  Dynamic CT can be used for image guided interventions, this technique is referred to as CT fluoroscopy.  CT fluoroscopy is routinely used for taking difficult biopsies. IAEA DR. KHALID IBRAHIM HUSSEIN - KKU DENTAL CT  CT scans of the jaw can be made with any regular CT scanner, but dedicated volume (cone beam) CT scanners for dental imaging are also available with flat panel detector.  These are designed for the patient to be seated.  Has a low contrast resolution, this means that soft tissues cannot be assessed appropriately in the reconstructed images. IAEA DR. KHALID IBRAHIM HUSSEIN - KKU CONTRAST ENHANCED CT  In contrast enhanced CT, contrast is artificially created between structures that would not be visible on non enhanced scans.  In CT angiography, iodine is administered during the CT scan intravenously to enhance the contrast between the vessels and the vessel wall.  In CTcolonography gas may be inflated through the rectum to enhance contrast between the colon and its surrounding tissues IAEA DR. KHALID IBRAHIM HUSSEIN - KKU SPECIAL APPLICATIONS  In CT for radiotherapy treatment planning the patient is scanned in the position that will be applied during the radiotherapy sessions.  Special wide bore scanners provide a gantry opening that is large enough to allow that a patient is scanned in such a position. They offer a large field of view. Standard field of view Extended field of view IAEA DR. KHALID IBRAHIM HUSSEIN - KKU SPECIAL APPLICATIONS  Dual energy CT imaging requires imaging of the volume of interest at two different (average photon) energies, which allow for better discrimination of certain tissues and pathology.  Dynamic CT imaging (4D CT): Is dynamic process in the volume of interest as a function of time. Can be used to visualize the movement of joints or the contrast enhancement in organs (perfusion or dynamic CT angiography). IAEA DR. KHALID IBRAHIM HUSSEIN - KKU COMPUTED TOMOGRAPHY IMAGE QUALITY Image quality  The CatPhan phantom that is widely used to evaluate image quality of CT scans.  To check the numerical value of Hounsfield units in the reconstructed image four large inserts in the periphery of the phantom represent air, -1000 HU, low density polyethylene, -100 HU, Acrylic, +115 HU, and Teflon, +990 HU, (the background is +90 HU).  Low contrast Acrylic inserts of different diameters around the center allow for exploring the effect of object size on low contrast IAEA detectability. DR. KHALID IBRAHIM HUSSEIN - KKU QUALITY CONTROL TEST FOR CT Daily Test performed by Technologist  Environmental Inspection  Tube warm-up  Air Calibration: done by automated scan taking several exposure and measured how much reaching the detector to calibrate the zero value DR. KHALID IBRAHIM HUSSEIN - KKU QUALITY CONTROL TEST FOR CT Standard Test from safety code 35 (Weekly tests)  CT # accuracy  Cross-field uniformity  Noise (Standards deviation) assisted in 40% of the size of the phantom. The uniformity module of Catphan 500 DR. KHALID IBRAHIM HUSSEIN - KKU QUALITY CONTROL TEST FOR CT Standard Test from safety code 35 (Monthly)  CT # Linearity : tested with phantom with different type of tissues with different density. The sensitometry module of Catphan 500 DR. KHALID IBRAHIM HUSSEIN - KKU QUALITY CONTROL TEST FOR CT  Spatial Resolution (high-contrast resolution) (Quarterly):  Directly, using a line pairs phantom (LP/cm)  Data analysis, Modulation Transfer Function (MTF) is graphical representation od system performance.  Can be described in-plane x-y direction and longitudinal in z-direction The Spatial Resolution and Slice thickness module of Catphan 500 DR. KHALID IBRAHIM HUSSEIN - KKU QUALITY CONTROL TEST FOR CT  Spatial resolution is limited by the acquisition geometry of the CT scanner, the reconstruction algorithm and the reconstructed slice thickness.  The performance of current 64-slice scanners with regard to spatial resolution, expressed as the full-width halfmaximum of the point spread function, is within the range 0.6-0.9 mm in all 3 dimensions DR. KHALID IBRAHIM HUSSEIN - KKU QUALITY CONTROL TEST FOR CT  Temporal Resolution:  The time required to complete an arc of rotation such that adequate photon energy has been received by the detector array & enabling reconstruction of an accurate representation of anatomical structures. 𝑅𝑜𝑡𝑎𝑡𝑖𝑜𝑛 𝑇𝑖𝑚𝑒 𝑇𝑒𝑚𝑝𝑜𝑟𝑎𝑙 𝑅𝑒𝑠𝑜𝑙𝑢𝑡𝑖𝑜𝑛 = 2 DR. KHALID IBRAHIM HUSSEIN - KKU QUALITY CONTROL TEST FOR CT  Contrast Resolution (CT low contrast detectability)  low-contrast resolution depends on tube voltage, beam filtration and the reconstruction algorithm.  Image noise is the main limitation for low-contrast resolution. Low contrast detectability module of Catphan 500 DR. KHALID IBRAHIM HUSSEIN - KKU I COMPUTED TOMOGRAPHY IMAGE QUALITY Artefacts Al  Artefacts can be related to data acquisition, image reconstruction, and the patient an  Data acquisition related examples D Ring artefact - malfunctioning of one or more detector elements ~ Unusable images from malfunctioning of the X-ray tube during the acquisition D The finite slice thickness leads to an averaging that is referred to as partial volume effect, S Beam Hardening artifacts: Change in the beam energy spectrum (Streaks and dark bands) IAEA DR. KHALID IBRAHIM HUSSEIN - KKU 2 COMPUTED TOMOGRAPHY IMAGE QUALITY ~ Artefacts  A ring artefact occurs in case of malfunctioning of one or more detector elements IAEA DR. KHALID IBRAHIM HUSSEIN - KKU Qa artifact 1. What are artifacts in computed tomography imaging, and what are the different types? Explain in detail one patient-related artifact and one reconstruction-related artifact. 2. What are metal artifacts and how can they be minimized? 3. What are beam hardening artifacts and how can they be minimized? 4. Discuss the partial volume effect and how it affects computed tomography image quality. 5. What are ring artifacts, and what causes them in computed tomography scans? 6. Explain how patient movement can lead to artifacts in computed tomography imaging and suggest ways to minimize these artifacts. 1) 2) 3) 4) 5) 6) Artifacts are errors in computed tomography (CT) imaging. There are two main types: patient-related artifacts (caused by patient movement) and reconstruction-related artifacts (errors during image reconstruction). An example of a patient-related artifact is motion artifact, which occurs when the patient moves during the scan. An example of a reconstruction-related artifact is partial volume artifact, which occurs when the CT slice thickness is greater than the size of the object being imaged.  Metal artifacts are streaks or distortions in CT images caused by the presence of metal objects. They can be minimized by using metal artifact reduction software, increasing the kVp of the scan, or using thinner slices.  Beam hardening artifacts are dark streaks or cupping artifacts in CT images caused by the hardening of the x-ray beam as it passes through the patient. They can be minimized by using beam hardening correction software, increasing the kVp of the scan, or using thinner slices.  The partial volume effect is the averaging of the CT numbers of different tissues within a voxel. This can lead to blurring of the image and loss of detail. The partial volume effect can be minimized by using thinner slices.  Ring artifacts are circular artifacts in CT images caused by a malfunctioning detector element.  Patient movement can lead to artifacts in CT imaging, such as motion artifacts and ghosting artifacts. These artifacts can be minimized by instructing the patient to remain still during the scan, using immobilization devices, or using motion correction software. 3 COMPUTED TOMOGRAPHY IMAGE QUALITY N ~ Artefacts  A partial volume effect: Is due to the finite slice thickness that leads to an averaging gray scale. IAEA DR. KHALID IBRAHIM HUSSEIN - KKU Y COMPUTED TOMOGRAPHY IMAGE QUALITY Q Hardening Artifacts  Strong attenuation of the X-ray beam by compact bone, ~ calcifications, or a metal object may lead to a beam hardening artifact.  Image shows typical streaks from an image of a large metal implant). Appear between two dense objects in the image. X IAEA Cupping artifacts Streaks and dark bands DR. KHALID IBRAHIM HUSSEIN - KKU 5 COMPUTED TOMOGRAPHY IMAGE QUALITY & Features to minimizing beam hardening  Filtration: Using bow-tie filters to pre-harden beam. & C Calibration Correction: Using uniform phantoms to correct Cupping  artifacts  Beam hardening correction software: Using Iterative correction & algorithm. IAEA Without Calibration With Calibration DR. KHALID IBRAHIM HUSSEIN - KKU b COMPUTED TOMOGRAPHY IMAGE QUALITY Artefacts byvolume effect when relatively thick slices are  Other reconstruction related artefacts include the partial reconstructed, the helical artefact (windmill patterns) in helical acquisitions, and the cone beam artefact (streaks). Windmill artefact IAEA DR. KHALID IBRAHIM HUSSEIN - KKU 7 COMPUTED TOMOGRAPHY IMAGE QUALITY Artefacts y  Patient related artefacts can sometimes be avoided by properly instructing the patient not to move during the scan and to maintain the breathhold during the entire scan, particularly during scans of the trunk. IAEA DR. KHALID IBRAHIM HUSSEIN - KKU QUANTITIES FOR CT DOSIMETRY Quantities for CT dosimetry  The irradiation conditions in CT are quite different from those in planar imaging and it is necessary to use special dosimetric quantities and techniques.  Measurements may be made free-in air or in-phantom  The dosimetric quantities for both are referred to as computed tomography kerma indices.  A pencil ionization chamber is generally used. DR. KHALID IBRAHIM HUSSEIN - KKU QUANTITIES FOR CT DOSIMETRY Terminology for CT dosimetry  Both types of measurement have in the past been expressed in terms of a ‘Computed Tomography Dose Index’ (CTDI) however, for measurements ‘in-phantom’ using an air kerma calibrated ionization chamber, the measured quantity is air kerma the absorbed dose to an air cavity within a phantom arises from a situation without secondary electron equilibrium and is difficult to measure  For these reasons, the terminology ‘Computed Tomography Kerma Index’ is used here for both free-in air and in-phantom measurements  All of the CT kerma indices used correspond directly with those previously referred to as CTDI related quantities DR. KHALID IBRAHIM HUSSEIN - KKU QUANTITIES FOR CT DOSIMETRY CT kerma index free-in-air (1 of 2)  The CT kerma index Ca,100, measured free-in-air for a single rotation of a CT scanner is The integral under the radiation dose along the z-axis from a single axial scan of width NT The integration range is positioned symmetrically about the volume scanned Ca,100 is usually expressed in units of mGy For in-phantom measurements the notation CPMMA,100 is used DR. KHALID IBRAHIM HUSSEIN - KKU QUANTITIES FOR CT DOSIMETRY CT kerma index free-in-air  From the equation it can be seen that the CT air kerma index is the height of a rectangular air kerma profile of width equal to the product of the number of sections, N, and the nominal section thickness, T, that has the same value as the line integral.  CTDI100 represents accumulated multiple scan dose from a series of contiguous irradiations DR. KHALID IBRAHIM HUSSEIN - KKU QUANTITIES FOR CT DOSIMETRY Weighted CT kerma index in phantom (1 of 2)  Unlike some areas of dosimetry, only two phantoms have found common application the standard head and body phantoms  The weighted CT kerma index, CW, combines values of C100,PMMA measured at the centre and periphery of these * phantoms ;, DR. KHALID IBRAHIM HUSSEIN - KKU QUANTITIES FOR CT DOSIMETRY Weighted CT kerma index in phantom (2 of 2) & O Center : 5 mg Y O & Pintry I = · &) ( + 3MGY 2, CPMMA,100,c is measured at the centre of the standard CT dosimetry # phantom CPMMA,100,p is the average of values measured at four positions around the periphery of the phantom  The weighted CT kerma index is an approximation to the - average air kerma in the volume of the phantom examine by a single rotation of the scanner # DR. KHALID IBRAHIM HUSSEIN - KKU Examples → → Hypothetical Measurements: Cn = 5 (PMM100 CPMMA,100,c & (center): 1.2 mGy (milligrays) CPMMA,100,p (periphery): 0.8 mGy (average of four peripheral measurements) Calculation: , c + 2 COMMA & - Cw = 1/3 (1.2 mGy + 2 * 0.8 mGy) Cw = 1/3 (1.2 mGy + 1.6 mGy) Cw = 1/3 (2.8 mGy) Cw = 0.93 mGy MA Example 2: CPMMA,100,c (center): 1.5 mGy CPMMA,100,p (periphery): 0.6 mGy Calculation: Cw = 1/3 (1.5 mGy + 2 * 0.6 mGy) Cw = 1/3 (1.5 mGy + 1.2 mGy) Cw = 1/3 (2.7 mGy) Cw = 0.9 mGy Example 3: CPMMA,100,c (center): 1.0 mGy CPMMA,100,p (periphery): 0.9 mGy Calculation: Cw = 1/3 (1.0 mGy + 2 * 0.9 mGy) Cw = 1/3 (1.0 mGy + 1.8 mGy) Cw = 1/3 (2.8 mGy) Cw = 0.93 mGy , 100, c = 1, 2 msY MGY , 100, p) QUANTITIES FOR CT DOSIMETRY X Volume averaged weighted CT kerma index in phantom  CVOL provides a volume average which takes account of the helical pitch or axial scan spacing - voln ast & of 2 > - CT > - ↓ (CPMMA , & e- waisted ?? Pitch 100, c + 2 Comma, 00, iP) -> where N is the number of simultaneously acquired tomographic sections T the nominal slice thickness & l is the distance moved by the patient couch per helical rotation or between consecutive scans for a series of axial scans p is the CT pitch factor (or pitch) for helical scanning T - - DR. KHALID IBRAHIM HUSSEIN - KKU Example : A CT scan is performed on a cylindrical phantom using the following parameters: Nominal beam width: 10 mm Pitch: 1.5 Measurements: ◦ CPMMA,100,c (center): 1.2 mGy ◦ CPMMA,100,p (periphery): 0.8 mGy (average of four peripheral measurements) ◦ CTDIw (weighted CT dose index): 10 mGy (this is an average over multiple rotations and slices) Calculations: 1. - Weighted CT Kerma Index (Cw): Cw = 1/3 (1.2 mGy + 2 * 0.8 mGy) Cw = 0.93 mGy 2- Volume CT Dose Index (CTDIvol): CTDIvol = CTDIw / Pitch CTDIvol = 10 mGy / 1.5 CTDIvol = 6.67 mGy QUANTITIES FOR CT DOSIMETRY CT pitch factor  p is the CT pitch factor (or pitch) for helical scanning where l is the distance moved by the patient couch per helical rotation or between consecutive scans for a series of axial scans N is the number of simultaneously acquired tomographic sections T the nominal slice thickness DR. KHALID IBRAHIM HUSSEIN - KKU Example: If a CT scan has a CTDIvol of 20 mGy and a scan length of 30 cm, the DLP would be: DLP = 20 mGy x 30 cm = 600 mGy·cm Absolutely! Let's illustrate with a comprehensive example involving DLP, CTDIvol, and Cw. Scenario: A patient undergoes a head CT scan. The following parameters and measurements are obtained: Scan Length: 16 cm Nominal Beam Width: 5 mm Pitch: 1.2 Measurements: ◦ CPMMA,100,c (center): 1.5 mGy ◦ CPMMA,100,p (periphery): 0.7 mGy (average of four peripheral measurements) ◦ CTDIw (weighted CT dose index): 12 mGy Calculations: T 1 Weighted CT Kerma Index (Cw): Cw = 1/3 (1.5 mGy + 2 * 0.7 mGy) Cw = 0.97 mGy 2 Volume CT Dose Index (CTDIvol): CTDIvol = CTDIw / Pitch CTDIvol = 12 mGy / 1.2 CTDIvol = 10 mGy 3 Dose-Length Product (DLP): DLP = CTDIvol * Scan Length DLP = 10 mGy * 16 cm DLP = 160 mGy*cm - BIBLIOGRAPHY  Buzug TM. Computed Tomography: From Photon Statistics to Modern Cone-Beam CT. ISBN 9783540394075. Springer. 2008  Hsieh J. Computed Tomography: Principles, Design, Artifacts, and Recent Advances. SPIE Press Book. ISBN: 9780819475336. Second edition, 2009  Kak AC, Slaney M. Principles of Computerized Tomographic Imaging, IEEE Press, 1988 (a free PDF copy is available at http://www.slaney.org/pct/ )  Kalender W. Computed Tomography: Fundamentals, System Technology, Image Quality, Applications. Publicis IAEA Corporate Publishing. 2005 DR. KHALID IBRAHIM HUSSEIN - KKU 8 , 5 May MGY 9.2 1. Given the values of CPMMA,100,c and CPMMA,100,p, calculate the weighted CT kerma index (CW). & 2. If N = 16, T = 5 mm, l = 40 mm, and p = 1.5, calculate the volume averaged weighted CT kerma index (CVOL). && T & 3. In a helical CT scan, if the patient couch moves 30 O mm per rotation and 10 slices are acquired simultaneously with a nominal slice&thickness of 2 mm, what is the CT pitch factor? & T & 4. DLP? I - 2 T - J D CW = - (COMMA , 100 1 + 2 COMMA + 2xa , ,5 -(8 - (8 , + 3 CW , 100 , 2) p) CVol = 18 , 4) , 8 , T CW = P I 30 max16x mym 126 9) = CrT Tox2 Youm 966 may = 8 , 966 E 17 , 232 mGY ? Koufa 16x5 Un N T