course: prh 499
course: prh 499

DAS data acquisition system
CT computed tomography
HU Hounsfield Units
DLP Dose Length Product
CTDI computed tomography dose index
ED Effective dose
MSCT Multislice ct
SSCT Single slice ct
ROI Region of interest
SD standard deviation
CS Contract scale
The radiology department of Princess Marina Hospital embodies six imaging modalities; ultrasound which involves the diagnostic use of ultrasound waves, computed radiography, digital radiography, mammography, fluoroscopy and computed tomography which use ionizing radiation. CT is a three dimensional imaging technique that uses a CT scanner. The CT scanner has a doughnut-shaped gantry that houses an x-ray tube and an array of detectors opposite the tube. During imaging the gantry rotates 360° as the CT bed on which the patient lies advances into it in the z-axis CITATION Kri17 l 1033 (Krikor Malajikian, 2017). It is usually recommended that patients are imaged for diagnostic purposes using all of the above methods as relevant except CT since it results in very high patient doses. In cases where the brain is to be imaged or when more detail than all the above modalities can give is needed, then CT is used as the imaging modality,
During the six week internship undergone in 2017 it was observed that there is only 1 CT scanner in Princess Marina Hospital. The scanner serves all the patients in the hospital who need CT images and those from many other government hospitals and clinics e.g. Scottish Livingstone hospital, Thamaga hospital and all the clinics in Gaborone. The knowledge that the quality of images produced by any imaging modality directly influences the visibility of structures on radiographs and the accuracy of diagnosis initiated interest in finding out just how good the CT scanner in Marina Hospital is performing since it serves so many patients. According toCITATION Kri14 l 1033 (Gulliksrud, 2014) International radiation protection authorities like ICRP and IAEA recommend acceptance tests and periodically QA-tests of CT scanners with respect to radiation dose and image quality. Consequently a 360° evaluation of the somatom scanner in Princess Marina Hospital was carried out in order to determine its performance in terms of image quality and patient doses(CTDI) using a catphan 500600 phantom. Then the images were processed using image owl. The image quality related tests that were carried out include, image uniformity and noise,CT number linearity and high and low contrast resolution.

Since every level of image quality in terms of one or more of the above factors comes with a radiation dose this project also evaluated the performance of the CT scanner in terms of radiation doses acquired during scanning. The performance of the scanner would be considered good if it shows CT number linearity, image uniformity and pixel noise, low and high contrast resolution equal or close to those recommended by image owl experts at acceptable doses.

Computed tomography (CT) is a computerized diagnostic imaging procedure used to create detailed cross-sectional images of internal organs, bones, soft tissue and blood vessels, representing in each pixel the local X-ray attenuation properties of the body. This procedure uses x-rays from a CT scanner’s xray tube to build the images which are reconstructed from measurements of attenuation coefficients of x-ray beams in the volume of subject studied. The CT scanner has a doughnut-shaped gantry that houses an x-ray tube and an array of detectors opposite the tube.

During imaging the gantry rotates 360° as the CT bed on which the patient lies advances into it in the z-axis CITATION Kri17 l 1033 (Krikor Malajikian, 2017). Instead of film, as in conventional x-ray imaging, CT scanners use special digital x-ray detectors, which pick up the x-rays as they leave the patient’s body from the opposite source and transmit them to the computer. Detector elements receive signals depending on the different attenuation coefficients µ in each cube along the distance they have to travel. The outgoing intensity I(x) of the beam of photons measured will depend on the location. In fact, I(x) is smaller where the body is more radio-opaque.Then, In order to obtain tomographic images of the patient from the data in “raw” scan, the computer uses complex mathematical algorithms for image reconstruction. The image of the section of the object irradiated by the X-ray, is reconstructed from a large number of measurements of attenuation coefficient. It gathers together all the data coming from the elementary volumes of material through the detectors. Using the computer, it presents the elementary surfaces of the reconstructed image from a projection of the data matrix reconstruction, the tone depending on the attenuation coefficients.

The X-ray attenuation mentioned severally above, within one pixel (also known as CT number) is expressed in Hounsfield units (HU) and air, fat, water and compact bone have the CT numbers Air = -1000, -60 to -120, 0and +1000 respectively.

CT number = 1000? (µ – µwater)/µwater (1)In this expression µ is the effective linear attenuation coefficient for the X-ray beam and air and water respectively have the following CT numbers: –1000 and 0 HU.

The CT image acquired is a digital image and consists of a square matrix of elements (pixel), each of which represents a voxel (volume element) of the tissue of the patient.

So, the signal transmitted by the detector is processed by the PC in the form of the digital information, the CT image reconstruction.
In  helical CT, the term pitch has two terminologies depending on whether single slice or multislice CT scanners are used . in single slice ct, The term detector pitch is used and is defined as table distance traveled in one 360° gantry rotation divided by beam collimation 2.

For example, if the table traveled 5 mm in one rotation and the beam collimation was 5 mm then pitch equals 5 mm / 5 mm = 1.0. 
Detector pitch affects both image quality and patient dose in that when pitch equals 1,  x-ray beams are contiguous for adjacent rotations, when greater than 1, x-ray beams are not contiguous for adjacent rotations, i.e. there are gaps in between the x-ray beams and tissue is not irradiated while when smaller than 1,  there is x-ray beam overlap; i.e. a volume of tissue is irradiated more than once per scan Which means a pitch greater than 1 results in decreased patient dose but also decreased image quality (through fewer projections obtained, resulting in lower SNR). A pitch less than 1 results in better image quality, but a higher patient dose. 
In Beam pitch is defined as table distance traveled in one 360° gantry rotation divided by total thickness of all simultaneously acquired slices 3.

During scanning the parameters that affect image quality are voltage and current .voltage (kV) e.g. 80-140 kV affects image quality in that the higher the voltage, better the penetration of x ray, but worse tissue contrast and larger dose . Electric current (mAs) of about 50-500 mAs affects image quality in that the higher the current, better the image quality (lower noise), but larger dose.

The cross-sectional images generated during a CT scan can be reformatted in multiple planes, and can even generate three-dimensional images which can be viewed on a computer monitor, printed on film or transferred to electronic media. And determine its size and location. CT is fast, painless, noninvasive and accurate. In emergency cases, it can reveal internal injuries and bleeding quickly enough to help save lives.

The generations of CT scanners are 1st generation CT Whereby the x-ray tube and single detector are connected and move together by translation and then rotation and the x-ray beam has linear (pencil-like) shape. A thin beam of X-rays was generated through the use of a collimator and a single detector element was used to measure the attenuated intensity. By translating this set-up, different positions were measured. After an entire set of parallel measurements had been acquired, the set-up was rotated to acquire the next parallel projection. In the 2nd generation CT the same type of movement is used and there are multiple detectors arranged in a row and a fan shaped x-ray beam instead of a linear shaped one. The 2nd generation CT scanners differed only slightly from that initial design in that a small number of measurement values could be obtained simultaneously. In Hounsfield’s first commercially available scanner a total of 180 projections were obtained in steps of 1° with 160 measurement values each. The acquisition of those 28.800 measurement values took five minutes. From that data an image of 80 × 80 pixels was reconstructed. With such a scanner, a head examination requiring six slices took about half an hour. Therefore,
The third generation has a fully rotating x-ray tube+detectors complex. Physicists were aiming at shortening the examination times. This was achieved with the introduction of the 3rd generation CT scanners: a 1D array of detector elements positioned on an arc covers the entire measurement field and acquires a complete ‘fan-beam’ projection (Figure 8). This not only avoided the slow translation movements, but also improved the efficiency of using the output of the X-ray tube. As figure 9 shows, a modern 3rd generation CT scanner is a machine consisting of a donut-shaped gantry with a big hole. Head, body, arms or legs have to be in the middle of the scanner to make a cross-sectional image. The patient is moved in and out on a motor-controlled table. The slice thickness is usually 0.5 to several mm and the spatial resolution (in the cross section) is roughly 1 mm at 512 × 512 pixels per slice. Within the ring of the CT scanner an X-ray tube is placed opposite a detector array with up to 1200 detecting elements, which receive the photons that went through the patient. If one measurement has been done this way, the source and detector rotate over a small angle (roughly 1°) and a new measurement is taken. The scanner repeats this procedure until a rotation of 180° has been reached. Then all thousands of measurements for reconstructing one slice have been done. The table on which the patient lies can then move a little further through the ring for measuring a new slice. In the fourth generation CT scanner only x ray tube rotates, detectors are stationary and this technology was later abandoned. CITATION Gia98 l 1033 (Douglas, 1998)RADIATION DOSE DESCRIPTORS IN CT
Before CT was introduced, planar radiography and fluoroscopy included the types of examinations in which radiation dose to the patient maximizes where the x-ray beam enters the surface of the skin. Consequently that made it reasonable to use radiation exposure to the entry surface (referred to as entrance skin exposure) as an indicator of radiation risk. The calculation of entrance skin exposure is straightforward, using measurements of in-air ionization chamber exposure at several x-ray tube kilovoltages covering the clinical range. Such measurements are usually expressed as exposure per milliampere second (mR/mAs, or more properly today, mGy air kerma/mAs).

However, during scanning of a CT slice, the x-ray beam enters from all directions at some point during the scan hence it is no longer clear where (on the surface of or inside the patient) the maximum dose occurs, nor is calculation of the dose at any point in or on the patient straightforward. For this reason radiation doses in CT are usually estimated using CTDI that is CT dose index. This is a standardized measure of radiation dose output of a particular CT scanner which allows the user to compare radiation output of different CT scanners or to estimate patient doses as shown below. CTDI has the following forms.

CTDI100 is a linear measure of dose distribution over a pencil ionization chamber and hence does not take into consideration the topographical variation of a human body and is therefore not in clinical use.

CTDIw is closer to the human dose profile as compared with the CTDI100
2/3 CTDI100 (periphery) +   1/3 CTDI100 (center)……………………………
CTDIvol is obtained by dividing CTDIw by pitch factor.

Another dose descriptor that is usually used is used is the dose-length product (DLP) which factors in the length of the scan to show overall dose output
DLP: CTDIvol x scan length…………………………………………………………
the CTDIvol or its derivative the DLP, as seen on consoles and outputted on the DICOM images, do not represent the actual absorbed or effective dose for the patient but should be taken as an index of radiation output by the system for comparison purposes.

If the AP and lateral dimensions of the patient are available, then the SSDE, that is the size-specific dose estimate can be used to estimate the absorbed dose.

Size specific dose estimate (SSDE) is a method of estimating CT radiation dose from a procedure that takes a patient’s size into account usually using 
CTDIvol and DLP . The exposures are the same regardless of patient size, but the size of the patients is a factor in the overall patient’s absorbed dose.

Coefficients were developed to help transform a patient’s exposure to absorbed dose if the AP and lateral dimensions of the patient are provided.

For example, to convert to SSDE the 2 formulas used are

where the size-specific conversion factor was defined as . Separate conversion factor tables were provided by Report 204 that depended on (a) the diameter of the PMMA phantom (either 16 or 32 cm) used to measure the generic CTDIvol and (b) the manner by which the patient size (X) was determined (Fig 2). The five methodologies suggested for determining patient size were as follows: (a)use of the anteroposterior (AP) parameter as a measurement of body thickness from anterior to posterior, (b) use of the lateral (LAT) parameter as a measurement of the body thickness from left to right, (c) the summation of the AP and lateral dimensions (sum = AP + LAT), (d) the calculation of the effective diameter of the patient (), and (e) the association of the patient’s age with their effective diameter as determined by International Commission on Radiation Units (ICRU) Report 74 (11). CITATION Sam12 l 1033 (Samuel L Brady, 2012)SSDE does not take the organs in the CT scan’s field of view into account, so it is not a measure of effective dose.

The effective dose is used to compare the stochastic risk of non-uniform exposure to radiation. Body tissues react differently to radiation and cancer-induction occurs at different rate of dose in different tissues. Hence, the effective dose is the risk of developing fatal cancer in the tissue in question. If the body is uniformly irradiated, the summed effective doses are equal to 1.

The effective dose is calculated by multiplying the equivalent dose (HT) by a tissue weighting factor(WT).
From all these formulars it camn be estab;lished that ctdi is directly proportional to effective dose.

Scanner Design Factors Affecting CT Radiation Dose
Both scanner design factors and clinical protocol factors affect radiation dose to the patient. Some design factors that affect the radiation dose required to achieve a particular image quality have been previously discussed (8). Those are the factors that determine dose efficiency (9).

The ability of a scanner to visualize low-contrast structures is inherently limited by image noise (quantum mottle). For any given radiation dose, maximum sensitivity requires capturing and using as many primary x-rays exiting the patient as possible. Dose efficiency, defined as the fraction of primary x-rays exiting the patient that contribute to the image, has 2 components: geometric efficiency (fraction of transmitted x-rays interacting with active detector areas) and absorption efficiency (fraction of actually-captured x-rays interacting with active detector areas). Geometric efficiency is reduced if some x-rays are absorbed before detection (e.g., in the detector housing) or if some x-rays do not enter active detector areas (e.g., by passing between detectors or striking inactive dividers between individual detectors). Absorption efficiency is reduced if some x-rays that enter the detectors are not absorbed.

Geometric efficiency for modern third-generation single-slice scanners is relatively high (?80%), with loss being due primarily to dead spaces between detector elements. Geometric efficiency is reduced in multislice CT relative to single-slice CT, because the separations between detector elements in the z-direction create more dead space and because more of the z-direction beam penumbra must be discarded (dose issues in multislice CT will be discussed in the third article of this series). Modern scanners generally use solid-state detectors, with absorption efficiency on the order of 99%.

Other design factors that may affect radiation dose include the distance of the x-ray tube from the isocenter (and thus from the patient), the design of the prepatient x-ray beam collimator, and the design of the bowtie filter and any other beam filtration.

Clinical Scanning Factors THAT AFFECT Radiation Dose IN CT
Radiation dose depends on tube current (amperage), slice scan time, and tube peak kilovoltage. As in radiography, tube current and slice scan time are taken together as mAs in relation to radiation dose and image quality. Increasing the mAs (by increasing tube current or slice scan time) increases the dose proportionally: 300 mAs deliver twice the dose of 150 mAs. Thus, CT radiation dose is often expressed as dose per mAs (or per 100 mAs).

Increasing peak kilovoltage (with all else held constant) also increases radiation dose, because the beam carries more energy. However, increasing peak kilovoltage significantly increases the intensity of the x-rays penetrating the patient to reach the detectors. Therefore, significantly lower mAs are needed to achieve similar image quality. Consequently, a higher peak kilovoltage does not necessarily mean an increased patient dose and, in fact, may allow the dose to be reduced.

CT, slice thickness, slice spacing, and helical pitch may affect dose as well. In single-slice CT with well-designed collimators, dose (as indicated by CTDI) is relatively independent of slice thickness for contiguous slices. Of course, the total length of the area scanned, as well as slice spacing, will determine how much total energy is deposited in the patient. For the same techniques, doses for helical scans with a pitch of 1.0 are equivalent to axial scans with contiguous slices. Pitches greater or less than 1 again affect CTDI values proportionally.

A relatively recent innovation allowing dose reduction in many cases is mA modulation. Before mA modulation came into use, a single mA value was specified (based on experience or the manufacturer recommendation) for the entire scan length, even though patient size or attenuation could change considerably along the scan length (e.g., compare attenuation through the thorax with that through the abdomen for a scan covering both areas). The result was often unnecessarily high mA values (and doses) for some slices, and a perhaps insufficient dose (and reduced image quality) for other slices. Using information from an initial scout view (a low-dose digital radiograph formed from a linear scan as the table moves through the gantry, with the x-ray tube stationary at, for example, 0° or 90°), the scan mA value is individually adjusted, depending on z-position, for each tube rotation. An enhanced version of mA modulation available on some scanners allows a mA adjustment not only for each rotation (z-position) but also as a function of angle during each rotation. Angle-dependent modulation is particularly valuable for anatomic regions in which a patient’s anteroposterior and lateral thicknesses are quite different (e.g., the pelvis). In such cases, the preselected mA value is often insufficient to provide adequate x-ray intensity at the detectors for lateral angles or may provide excessive intensity at the detectors for anteroposterior/posteroanterior angles. Angular mA modulation optimizes mA selection for each angle to provide the least radiation dose for the required level of image quality

Fundamentally, image quality in CT, as in all medical imaging, depends on 4 basic factors: image contrast, spatial resolution, image noise, and artifacts. Depending on the diagnostic task, these factors interact to determine sensitivity (the ability to perceive low-contrast structures) and the visibility of details.

CT Image Contrast
Low Contrast resolution OR CT IMAGE CONTRAST IS THE the ability to distinguish between differences in intensity in an image. It measures the ability of the CT scanner to distinguish relatively large objects that differ only slightly in density from background. On the other hand high contrast resolution CITATION DSS06 l 1033 (D.S Sharma, 2006) or spatial resolution was defined as the ability of the imaging modality to differentiate two objects which are closely located as individual objects and not just one bigger object.
CT image contrast depends on subject contrast and display contrast. Because CT display contrast is arbitrary (depending only on the window level and width selected), it will not be discussed further.

As in radiography, CT subject contrast is determined by differential attenuation: that is, differences in x-ray attenuation by absorption or scattering in different types of tissue and thus resulting in differences in the intensity of the x-rays ultimately reaching the detectors. Because of the high peak kilovoltage and relatively high beam filtration (beam hardness) used in CT, the x-ray/tissue interactions (except in bone) are overwhelmingly Compton-scattering events. Differential attenuation for Compton scatter arises from differences in tissue electron density (electrons/cm3), which in turn are due primarily to differences in physical density14. Thus, subject soft-tissue contrast in CT comes mainly from differences in physical density. That the small differences in soft-tissue density can be visualized on CT is due to the nature of the image (a 2-dimensional image of a 2-dimensional slice), the ability to map small attenuation differences to large differences in gray level by windowing, the near-complete elimination of scatter, and the use of a sufficient x-ray intensity.

Related to CT image contrast is the CT contrast scale. We recall that CT numbers are derived from voxel attenuation coefficients calculated during image reconstruction using the following relationship:Eq. 7where ?p and ?w are linear attenuation coefficients for a given voxel and for water (?w is determined from calibration scans). Because CT number is a linear function of ?p, a graph of expected CT numbers for materials with known attenuation coefficients should be linear over the clinical CT-number range (e.g., ?1,000 to +1,000). For evaluation of the contrast scale of a scanner, various CT test phantoms are available that contain materials designed to provide certain CT numbers (e.g., the CT numbers for water, fat, soft tissue, bone, and air)
High contrast resolution CITATION DSS06 l 1033 (D.S Sharma, 2006) or spatial resolution is the ability of the imaging modality to differentiate two objects which are closely located as individual objects and not just one bigger object. Resolution is the measure of how far two objects must be apart before they can be seen as separate details in the image. For two objects to be seen as separate the detectors must be able to identify a gap between them.Resolution is measured in line pairs per centimeter (lp/cm) i.e. the number of line pairs that can be imaged as separate structures within one centimeter.

Spatial resolution in CT, as in other modalities, is the ability to distinguish small, closely spaced objects on an image. A common test is an evaluation of limiting resolution, performed using line-pair test patterns. CT phantom line-pair patterns consist of bars of acrylic (or some denser plastic) separated by spaces containing a material that is less attenuating. The widths of the bars and spaces are equal and typically range from about 0.05 or smaller to 0.5 cm. Examples are shown in Figure 8. Bars of lead or other dense materials would cause severe artifacts on CT images and thus are not used. Resolving a line-pair test pattern requires that each bar and space be separately visible on the image. Each bar plus adjacent space is referred to as a line-pair. Rather than specifying bar width, bar pattern sizes are usually described by a spatial frequency in line-pairs per centimeter, defined as follows, where bar width is in centimeters:Eq. 8For example, a pattern with 0.1-cm bars and spaces has a spatial frequency of 1/(2 × 0.1), or 5 line-pairs per centimeter. In radiographic imaging, the x-ray tube focal-spot size and blur occurring in the image receptor are the primary causes of reduced resolution. Although focal-spot size does affect CT spatial resolution, CT resolution is generally limited by the size of the detector measurements (referred to as the aperture size) and by the spacing of detector measurements used to reconstruct the image. This concept, called sampling, is illustrated in Figure 9. In Figure 9A, consider a scan of a phantom containing a hypothetical test pattern: for example, 5 line-pairs per centimeter (0.1-cm bars) and detectors that are twice as wide (say, 0.2 cm and spaced by 0.2 mm). The aperture is approximately equal to the detector width—a width that in this case is clearly too large to resolve the smaller bars: All measurements (shown at the bottom of the figure as view data) include attenuation corresponding to half bar and half space. Suppose, instead, that the detector size is reduced to the size of the bars but the spacing is kept at 0.2 cm (5 samples per centimeter; Fig. 9B). Now, even though the aperture is sufficiently small, the bars still are not resolved because the samples are too far apart (in this case, the spaces between bars are missed). In addition to a small aperture, closely spaced measurements are required for good resolution (Fig. 9C). The general rule, known as the Nyquist criterion, states that resolving N line-pairs per centimeter requires measuring at least 2 × N samples per centimeter. For example, resolving 5 line-pairs per centimeter (0.010-cm bars) requires at least 10 measurements per centimeter. Some factors of scanner design associated with sampling were discussed earlier (8).

 There are two types of resolution in CT scanning:
Transaxial resolution (7 lp/cm)
Axially across the patient
Z-sensitivity (0.5 – 10 mm)
Along the length of the patient in the z-direction
Transaxial resolution
The minimum transaxial resolution is determined by the actual detector size, however it is often quoted as the “effective detector width” at the isocenter of the scanner (center of the bore of the scanner). The “effective detector width” and the actual detector size are slightly different due to the divergence of the beam. The smaller the “effective detector width” the higher the resolution.

The transaxial resolution is affected by scanner (hardware) factors or scan and reconstruction parameters.

Scanner factors
1. Focal spot
Smaller focal spots give higher resolution, but the max mA is limited to prevent damage to the anode.

There are usually two available focal spot sizes on CT scanners, for example:
Fine = 0.7 mm
Broad = 1.2 mm
Flying focal spot: the position of the focal spot is rapidly altered in the transaxial plane and/or the Z-axis. Each focal spot position increases the number of projections sampled and improves spatial resolution. For example, if the position of the focal spot moves in the X-Y plane, then the in-plane resolution increases.

Focus-detector distance (FDD)
Focus-isocentre distance (FID)
2. Detector size
Smaller detectors give higher resolution but more detectors within an area also means more partitions (dead space) and a reduced overall detection efficiency.
3. Detector design properties
Quarter ray detector offset: the Center of the detector array is offset from the center of rotation by one quarter the width of an individual detector. As the gantry rotates to 180° the centre of the detector array is now offset by half the width of a detector giving an interleaved sampling of the patient.

Scan parameters
1. Number of projections
Larger number of projections gives finer resolution (up to a point).

2. Reconstruction filter
Higher resolution or “sharp” kernels (e.g. bone reconstruction) have better spatial resolution than soft kernels (e.g. soft tissue reconstruction).

However, higher resolution kernels do not average high spatial frequency signals and therefore produce more noise.

3. Pixel size
The pixel size (d) in mm is give by the equation:
d = FOV/n
FOV = field of view (mm)n = image matrix size
The highest spatial frequency that can be obtained (fmax) is called the Nyquist limit and is given by:
fmax = 1/2d
From this equation you can see that the higher the pixel size, the lower the maximum spatial frequency.

To improve spatial frequency we can:
Reduce the field of view (smaller FOV = smaller pixel size as seen in the first equation). We can do this retrospectively by a targeted reconstruction of the original data into a small field of view.

Increase the matrix size (larger n = small pixel size as seen in the first equation)
Z-sensitivity refers to the effective imaged slice width.

Factors affecting z-sensitivity
1. Detector slice thickness
The wider (in the z-axis) the detector row, the lower the resolution
2. Overlapping samples
Acquiring the data using overlapping slices can improve Z-sensitivity. This is achieved by using a low spiral pitch e.g. pitch ;1.

3. Focal spot
A fine focal spot improves the z-sensitivity
Importance of slice thickness
1. Noise
The thinner the slice the better the resolution BUT the worse the noise
2. Partial volume effect
Thicker slices increase the partial volume effects
3. Isotropic scanning
Thin slices allow isotropic scanning, i.e. the pixels in the axial and the z-axis are the same size (cubes). The advantages of this are:
Reduced partial volume effect
Better multi-planar reformatting
Improved volume rendering e.g. displaying 3D representations of the data (e.g. cardiac imaging, vascular imaging, CT colonography etc)
Other CT Factors Affecting Spatial Resolution
Focal-spot size does affect CT resolution but to a lesser extent than in radiography. Motion can also introduce blurring, although the more important effect of motion is the potential creation of artifacts. Two additional factors—one potential and one common—are matrix size and reconstruction filter.

The displayed spatial resolution may be affected by reconstruction or by display pixels that are too large. A fundamental limitation is the size and spacing of detector measurements. Suppose that it ought to be possible, on the basis of sampling, to resolve 10 line-pairs per centimeter (0.05-cm bars). Now suppose that a 512 × 512 pixel matrix is reconstructed representing a 50-cm scan circle (i.e., the image represents a 50-cm-diameter area, which might be the case for a scan of a large body). The size of the pixels is approximately 0.1 × 0.1 cm (i.e., 50 cm/512 pixels), which is too large to resolve the 0.05-cm bars. It is often possible, if raw data are not yet overwritten, to reconstruct an image over a smaller circle, say 25 cm, to yield smaller pixels and higher resolution. Similarly, if the display pixels are larger than the reconstruction matrix pixels (uncommon in modern scanners), full resolution will not be displayed. If so, a graphic zoom (a function available on most scanners) will provide better resolution.

Although pixel size may affect resolution (e.g., for large scan circles), a reconstruction filter always affects resolution—often dramatically. A reconstruction filter is applied during filtered backprojection reconstruction to remove the blur from images (8). Usually, however, the filter is deliberately chosen to produce somewhat blurry images. The reason is that overly sharp CT images are usually too corrupted by image noise for most diagnostic tasks. A blurry filter also blurs noise and thus produces better diagnostic quality. For imaging tasks requiring more detail (e.g., to view bone), the operator may optionally select a sharper filter when setting the parameters for the scan. Commonly used filters are designed as compromises between acceptable spatial resolution and an acceptable level of noise. For example, a standard filter may produce images with a maximum resolution of, say, 6 line-pairs per centimeter, whereas a bone filter may image with a resolution of 10 line-pairs per millimeter or better. The test phantom is shown reconstructed with both a standard filter (Fig. 8A) and a high-resolution (bone) filter Fig. 8B
Ideally, a homogeneous material e.g. a water filled object, the CT numbers of that object have to be consistent throughout the material (image uniformity) across the scan field. However in CT imaging there is still variation in Hounsfield units about a mean. Image noise is usually measured via the signal to noise ratio (SNR); comparing the level of desired signal (photons) to the level of background noise (pixels deviating from normal). The higher the ratio, the less noise is present on the image. Noise on a cross-sectional image will equal a decrease in the picture quality and hinder the contrast resolutionCITATION DrD l 1033 (Murphy, 2005). The noise reduces image quality in that it degrades low contrast resolution and introduces uncertainty in the Hounsfield units of the images.
The three sources of noise include Quantum noise, Electronic noise and Noise introduced by the reconstruction process e.g. back projection but the most dominant of them is due to the fact that photon registration by the detectors is a stochastic process whereby the number of photons detected vary randomly about a mean value and that variation is the noise
Noise is usually measured using any uniform region of the image e.g. with a water phantom or the homogeneous module in a catphan phantom using regions of interest (ROIs) on a scan of the uniform phantom. A statistical ROI function (available on most CT scanners) allows users to place a rectangular or oval ROI on the image, within which is calculated the average and standard deviation (SD) of the CT numbers for the enclosed pixels. The SD indicates the magnitude of random fluctuations in the CT number and thus is related to noise: The larger the SD, the higher the image noise. A more complete discussion of image noise is presented in Because noise is the most bothersome when one is viewing low-contrast soft-tissue structures, an important test of scanner performance is how well low-contrast test objects are seen in the presence of typical noise levels. is an example of a low-contrast test phantom, consisting of groups of rods embedded in material producing approximately 0.6% subject contrast (i.e., a nominal CT-number difference of 6 between the rods and the background). The rod groups range in diameter from 6 to 2 mm. In this example, the 5-mm rods are visible, whereas the smaller ones are lost in the noise.

Then the standard deviation of the CT number in a selected region-of-interest gives the mean noise measurement.

The noise in the final image is given by:
Noise (standard deviation) ? 1/? (no. of photons)
For a well-designed CT scanner, image noise (quantum mottle) should be statistical: that is, due to random variations in detected x-ray intensity (electronic and other noise sources should be minimal in comparison). Quantitatively, these statistical fluctuations are described by the Poisson distribution, which states that the size of random variations (referred to as the SD) associated with measuring N x-rays is given by the square root of N. For example, if we detect 10,000 x-rays and then repeat this measurement several times, the measurements will not be exactly 10,000 each time but will fluctuate around an average or mean value of 10,000. The size of the random fluctuations will be on the order of 100 (the square root of 10,000). We would thus say that our measurement was 10,000 ± 100.

From this equation, it is evident that increasing the number of photons reduces the amount of noise and, therefore, anything that increases the number of photons (increases the photon flux) will reduce the noise. If we double the number of photons we will reduce the noise by ?2 (i.e. increasing the number of photons by a factor of 4 will halve the noise).

In radiography, image noise is related to the numbers of x-ray photons contributing to each small area of the image (e.g., to each pixel of a direct digital radiograph). In CT, x-rays contribute to detector measurements and not to individual pixels. CT image noise is thus associated with the number of x-rays contributing to each detector measurement. To understand how CT technique affects noise, one should imagine how each factor in the technique affects the number of detected x-rays. Examples are as follows:
X-ray tube amperage: Changing the mA value changes the beam intensity—and thus the number of x-rays—proportionally. For example, doubling the mA value will double the beam intensity and the number of x-rays detected by each measurement.

Scan (rotation) time: Changing the scan time changes the duration of each measurement—and thus the number of detected x-rays—proportionally. Because amperage and scan time similarly affect noise and patient dose, they are usually considered together as mA × s, or mAs.

Slice thickness: Changing the thickness changes the beam width entering each detector—and thus the number of detected x-rays—approximately proportionally. For example, compared with a slice thickness of 5 mm, a thickness of 10 mm approximately doubles the number of x-rays entering each detector. The number of photons available to generate an image has a linear relationship to the slice thickness. The thicker the slice, the more photons available, the more photons available; the better the SNR, this isn’t without a tradeoff, increasing the slice thickness will decrease the spatial resolution in the z-axis.

Peak kilovoltage: Increasing the peak kilovoltage increases the number of x-rays penetrating the patient and reaching the detectors. Thus, increasing the kilovoltage reduces image noise but can (slightly) reduce subject contrast as well. CITATION drS11 l 1033 (Abdulla, 2011)patient size
Larger patients will absorb more radiation than smaller ones, meaning fewer photons will reach the detector hence reducing the signal to noise ratio.

Although not affecting the numbers of detected x-rays, a reconstruction filter profoundly affects the appearance of noise in the image: Smooth filters blur the noise, reducing its visual impact, whereas sharp filters enhance the noise. In images of soft tissue, noise is generally more interfering than blur, and smoother filters are preferred. In images of structures with edges and small details, such as bone, blur is generally more interfering than noise, and sharper filters are preferred. For comparison, Figure 10 shows examples of noise in scans of uniform phantoms using standard and higher-resolution (bone) filters and with standard and very low values for mAs
Doubling the number of photons can be achieved by:
Doubling the tube current Doubling the rotation time (s)
Doubling the slice thickness (mm)
Increasing the tube kilovoltage (kV) also increases the photon flux but it is not directly proportional (output is approximately ? kV2).

Linearity is a property of a system, characterized by output that is directly proportional to the input.In computed tomography (CT), linearity describes the amount to which the CT numberof a material is exactly proportional to the density of this material (in Hounsfield units). This accuracy between the linear attenuation coefficient and the CT number is also utilized to describe the performance of a CT scanner.The linearity of a gamma camera is a measure of the geometrical correctness of the imagesCATPHAN PHANTOM
Initial phantom positioning
The Catphan® phantom is positioned in the CT scanner by mounting it on the case. Place the phantom case on the gantry end of the table with the box hinges away from the gantry. It is best to place the box directly on the table and not on the table pads. Open the box, rotating the lid back 180°. If you are using an annulus, additional weight will need to be placed in the box to counterweigh the phantom. The patient straps can be used for additional stability. Remove the phantom from the box and hang the Catphan® from the gantry end of the box. Make sure the box is stable with the weight of the phantom and is adequately counterweighed to prevent tipping. Use the level and adjusting thumb screws to level the Catphan®. Once the phantom is level, slide the phantom along the end of the box to align the section center dots on the top of the phantom with the x axis alignment light. Use the table height and indexing drives to center the first section’s (CTP401 or CTP404, Slice Geometry) alignment dots on the side and top of the phantom with the scanner alignment lights.

The z axis scan alignment position can be selected from the localizer scan, by centering the slice at the intersection of the crossed wire image created by the slice width ramps. Scan the first section (CTP401 or CTP404) and check the image for proper alignment as illustrated in the Phantom position verification section.

Incremental phantom module positioning
The Catphan® phantoms are designed so all test sections can be located by precisely indexing the table from the center of section 1 (CTP401 or CTP404) to the center of each subsequent test module. This design eliminates the need to remount the phantom once the position of section 1 (CTP401 or CTP404) has been verified. The indexing distances from section 1 are listed below. Additional illustrations on the preceeding page show the test modules and their index spacing. Phantom position and alignment verification is described on the next page.

Catphan® 500600 test module locations:
Module Distance from section 1 center
CTP528, 21 line pair high resolution 30mm
CTP528. Point source 40mm
CTP515, Subslice and supra-slice low contrast 70mm
CTP486, Solid image uniformity module 110mm
Phantom position verification By evaluating the scan image of section 1 (CTP401 or CTP404) the phantom’s position and alignment can be verified. The section contains 4 wire ramps which rise at 23° angles from the base to the top of the module. The schematic sketches below indicate how the ramp images change if the scan center is above or below the z axis center of the test module. The use of the scanner’s grid image function may assist in evaluation of phantom position.

Correct alignment
In this image the x, y symmetry of the centered ramp images indicates proper phantom alignment
Sensitometry (CT number linearity)
Four or seven high contrast sensitometric targets surround the wire slice thickness ramps. Three are made from the commercial plastics: Teflon, acrylic and low density polyethylene (LDPE). The fourth is air. These targets range from approximately +1000 H to -1000 H.

The monitoring of sensitometry target values over time and can provide valuable information, indicating changes in scanner performance.

Linear attenuation coefficient µ units cm-1
KEV Teflon Delrin Acrylic Polystryrene Water LDPE PMP Air
40 0.556 0.327 0.277 0.229 0.240 0.209 0.189 0
50 447 0.283 0.244 0.209 0.208 0.191 0.173 0
60 0.395 0.260 0.227 0.196 0.192 0.181 0.164 0
62 0.386 0.256 0.224 0.194 0.190 0.179 0.162 0
64 0.380 0.253 0.221 0.192 0.188 0.178 0.160 0
66 0.374 0.251 0.219 0.191 0.186 0.17 0.160 0
68 0.370 0.248 0.217 0.189 0.184 0.175 0.158 0
70 0.363 0.245 0.215 0.188 0.182 0.174 0.157 0
72 0.359 0.243 0.214 0.186 0.181 0.172 0.155 0
74 0.355 0.240 0.211 0.185 0.179 0.171 0.155 0
76 0.351 0.23 0.210 0.184 0.178 0.170 0.154 0
78 0.346 0.236 0.208 0.183 0.177 0.168 0.152 0
80 0.342 0.234 0.207 0.180 0.175 0.167 0.151 0
90 0.328 0.225 0.199 0.175 0.170 0.163 0.147 0
CTP528 High resolution module with 21 line pair per cm gauge and point source

This section has a 1 through 21 line pair per centimeter high resolution test gauge and two impulse sources (beads) which are cast into a uniform material. The beads are positioned along the y axis 20mm above or below the phantom’s center and 2.5 and 10mm past the center of the gauge in the z direction. On older CTP528 modules the bead is aligned in the z axis with the gauge.

21 Line pair per centimeter high resolution gauge
The 21 line pair/cm gauge has resolution tests for visual evaluation of high resolution
ranging from 1 through 21 line pair/cm. The gauge accuracy is ± 0.5 line pair at the 21
line pair test and even better at lower line pair tests.

The gauge is cut from 2mm thick aluminum sheets and cast into epoxy. Depending on the
choice of slice thickness, the contrast levels will vary due to volume averaging.

Line Pair/cm Gap Size Line Pair/cm Gap Size

CTP515 low contrast module with supra-slice and subslice contrast targets
The low contrast targets have the following diameters and contrasts:
Supra-slice target diameters Subslice target diameters
2.0mm 3.0mm
3.0mm 5.0mm
4.0mm 7.0mm
5.0mm 9.0mm
Nominal target contrast levels
Since the target contrasts are nominal, the actual target contrasts need to be determined
before testing specific contrast performance specifications. The actual contrast levels are
measured by making region of interest measurements over the larger target, and in the
local background area. To determine actual contrast levels, average the measurements
made from several scans. It is important to measure the background area adjacent to the
measured target because “cupping” and “capping” effects cause variation of CT numbers
from one scan region to another. Position the region of interest to avoid the target
edges. The region of interest should be at least 4 x 4 pixels in diameter. Because low
contrast measurements are “noisy” it is advisable to calculate the average of the multiple
measurements made from several scans. Carefully monitor the mAs setting because
the photon flux will improve with increased x-ray exposure. Use the size of the targets
visualized under various noise levels to estimate information on contrast detail curves.

CTP486 Image uniformity module

The image uniformity module is cast from a uniform material. The material’s CT number is designed to be within 2% (20H) of water’s density at standard scanning protocols. The typically recorded CT numbers range from 5H to 18H. This module is used for measurements of spatial uniformity, mean CT number and noise value. The precision of a CT system is evaluated by the measurement of the mean value and the corresponding standard deviations in CT numbers within a region of interest (ROI). These measurements are taken from different locations within the scan field.

The mean CT number and standard deviation of a large number of points, (say 1000 for example) in a given ROI of the scan, is determined for central and peripheral locations within the scan image for each type of scanning protocol. Inspect the data for changes from previous scans and for correlation between neighboring slices.

Measure spatial uniformity by scanning the uniformity section. Observe the trends
above and below the central mean value of a CT number profile for one or several rows or
columns of pixels as shown above.

Select a profile which runs from one side of the uniformity module to the opposite side.

Due to scanner boundary effects, typical profiles start 2cm from the edge of the test

Integral uniformity may be measured by determining the minimum and maximum CT
In 1999 a study was carried out to characterize the performance of a recently introduced multi?slice CT scanner in comparison to a single?slice scanner from the same manufacturer. The study involved the assessment of performance parameters such as: radiation and slice sensitivity profiles, low?contrast and limiting spatial resolution, image uniformity and noise, CT number and geometric accuracy, and dose. The multi?slice system was tested in axial (1, 2, or 4 images per gantry rotation) and HQ (Pitch = 0.75) and HS (Pitch = 1.5) helical was found out that the axial and HQ?helical modes of the multi?slice system provided excellent image quality in terms of the above parameters and a substantial reduction in exam time and tube loading, although at varying degrees of increased dose relative to the single?slice scanner.CITATION Cyt99 l 1033 (Cy.thia H.McCollough, 1999) In axial modes imaging planes that cut the body into superior and inferior are used. In the helical modes the gantry rotates 360 degrees completing a slice then the CT bed moves into it and another slice is taken throughout the specified section of the body.

In 1999 a study was carried out to characterize the performance of a recently introduced multi?slice CT scanner in comparison to a single?slice scanner from the same manufacturer. The study involved the assessment of performance parameters such as: radiation and slice sensitivity profiles, low?contrast and limiting spatial resolution, image uniformity and noise, CT number and geometric accuracy, and dose. The multi?slice system was tested in axial (1, 2, or 4 images per gantry rotation) and HQ (Pitch = 0.75) and HS (Pitch = 1.5) helical was found out that the axial and HQ?helical modes of the multi?slice system provided excellent image quality in terms of the above parameters and a substantial reduction in exam time and tube loading, although at varying degrees of increased dose relative to the single?slice scanner.CITATION Cyt99 l 1033 (Cy.thia H.McCollough, 1999) In axial modes imaging planes that cut the body into superior and inferior are used. In the helical modes the gantry rotates 360 degrees completing a slice then the CT bed moves into it and another slice is taken throughout the specified section of the body.

In 2006 D.S Sharma, K. K Sanu and others carried out a study in which comprehensive tests were done on a single slice CT scanner using in-house fabricated phantoms to validate the auto-performance test (APT) results. Electromechanical, image quality related and radiation safety related tests were carried out and the results for electromechanical parameters were found within the specified limits. The percentage noise obtained by APT was 1.32% while the independently measured value was 0.38%. Observed contrast resolutions by independent method at 0.78% and 12% contrast difference were 4 mm and 1.25 mm (= 4 lp/cm) respectively. However, high contrast resolution (limiting spatial resolution) by APT at 50, 10 and 2% MTF levels were 9, 12.5 and 14.1 lp/cm respectively. Calculated and measured CT numbers of water, air, teflon, acrylic, polystyrene and polypropylene differed by 0 to 24 HU, while this difference was 46 and 94 HU in case of nylon and bakelite respectively. The contrast scale determined using CT linearity phantom was 1.998×10?4 cm?1/CT number. CT dose index (CTDI) and weighted CTDI (CTDIw) measured at different kVp for standard head and body phantoms were smaller than manufacturer-specified and system-calculated values and were found within the manufacturer-specified limit of ± 20%. Measured CTDIs on surface (head: 3.6 cGy and body: 2.6 cGy) and at the center (3.3 cGy, head; and 1.2 cGy, body) were comparable to reported values of other similar CT scanners and were also within the industry-quoted CTDI range. CITATION DSS06 l 1033 (D.S Sharma, 2006)In the above study in-house fabricated phantoms were used for different tests instead of the catphan cylindrically shaped phantom that was used in this current project. The in-house fabricated phantoms are disc like in shape and are similar to those contained within the catphan phantom so the catphan phantom and the in-house fabricated phantoms combined, both give a sufficient report on image quality of the scanner. The catphan phantom was not used because the above study used a single slice CT scanner so the individual in-house fabricated phantoms were the best to facilitate that since it will take only one slice of each small phantom each time. With the multislice scanner catphan phantom is the best option.

Another study was carried out in the University of California in 2017 to evaluate the feasibility of breast computed tomography in terms of dose and image quality that is to evaluate the performance of a CT scanner in breast imaging. The tests carried out included signal signal-to-noise ratio in the images, noise and doses incurred. During the research it was found out that the maximum dose at mammography assessed in 1-mm3 voxels was far higher (20.0 mGy) than that at breast CT (5.4 mGy) for a typical 5-cm 50% glandular breast and CT images of an 8-cm cadaveric breast (AGD, 6.3 mGy) were subjectively superior to digital mammograms (10.1 mGy) of the same specimen. Consequently it was concluded that the potential of high signal-to-noise ratio images with low anatomic noise, which are obtainable at dose levels comparable to those for mammography, suggests that dedicated breast CT should be studied further for its potential in breast cancer screening and diagnosis. CITATION Joh17 l 1033 (John M. Boone, 2017)What is common with all of the above sources is that in evaluating the performance of the CT scanners in those places, the image quality tests included image uniformity and noise, high and low contrast resolution and CT number linearity in most of the studies and all of them included contrast and resolution. This is because image contrast and spatial resolution directly affect the ease with which structures are visualized on an image and the accuracy of diagnosis given consequently. If a CT scanner is able to differentiate signal intensities and reconstructs images in which there is much visible difference between a structure of interest and the background, then it will be easy for the radiologist to spot it and give an accurate report on it. The same applies to if the CT scanner is able to resolve or discern and depict very closely located objects as individuals and not as one bigger object.

In addition, as they evaluated the performance of their scanners, the above researchers included radiation safety tests of the scanners, the patient doses incurred during scanning. This is because when using an imaging modality that uses ionizing radiation, image quality and patient doses are inseparable. Any level of image quality produces a particular amount of radiation dose so the best option would be just to quantify both of the two factors and work at achieving a balance between them; that is trying to use the lowest possible doses that can give a visible-enough image for the best possible diagnosis.
Machinery and equipment used in this research include the somatom CT scanner and the Catphan 500600 phantom in Princess Marina Hospital. The Catphan 500600 phantom is a cylindrically-shaped The 20 cm long phantom and consists of four disc-shaped, 4cm thick inner phantoms and can measure slice width, pixel size, sensitometry, high contrast resolution, low contrast resolution and image uniformity. However this research focused on low and high contrast resolutions, sensitometry, image uniformity and percentage noise since they are most important in describing image quality CITATION Kri14 l 1033 (Gulliksrud, 2014).Image owl system will also be used to process the acquired images.

The activities included
1. Taking CT images of the Catphan phantom
First the containment case of the phantom was placed on the gantry end of the CT bed with the box hinges away from the gantry. Then the box was opened rotating the lid back 180°. Then the phantom was removed from the box and the hung from the gantry end of the box. (Making sure the box was stable with the weight of the phantom and adequately counterweighed to prevent tipping).

After that the horizontally lying phantom was aligned by moving the CT bed up and down in a way that resulted in the alignment lights of the scanner coinciding with the dots on top and on the sides of the phantom.

Then CT images were taken across the whole phantom using the exposure factors e.g. KVp and mA selected by the control console in the hospital. This took place in the hospital with the help of a radiographer.

2. Estimating the results by observation and calculations
Results were then estimated by observation and calculations with formulas from the Catphan 500600 manual and as guided by it to have an idea of what to expect from the image processor for more accurate results.

High contrast resolution
Using the spatial resolution module image which has resolution of 1 to 21 line pairs per cm the spatial resolution of the CT scanner was estimated by observation. The highest number lpcm that was resolved in the image will be taken to be the spatial resolution of the scanner.

Low contrast resolution Low-contrast resolution measures the ability of CT scanner to distinguish relatively large objects that differ only slightly in density from background and was be determined using the low-contrast slice image.

C. CT-number linearity
Here the sensitometry slice image was used. This slice has different materials such as acrylic, Teflon, LDPE embedded in plastic material. These materials have different attenuation factors at different kVp and different CT numbers ranging from +1000 to -1000.using the linear attenuation factors provided with in the manual for different tube voltages, and CT numbers calculated from them for each of the materials, a plot of CT number vs linear attenuation factor was drawn and a line of best fit was drawn to see if the trend will show CT number to be directly proportional to the linear attenuation factors for all those materials. If the line of best fit took the form of a straight line with a positive gradient(y=mx+c) then the scanner would estimated to have CT number linearity.

D. image uniformity and pixel noise: 
The image of the homogeneous image uniformity slice was used here. Ct numbers was determined at different positions in the image e.g., the 12, 6, 3 and 9 clock face positions as well as at the center to see if the CT numbers will be the same since the material is homogeneous.

% of noise = (? × CS × 100)/?w 
Where ? = standard deviation of CT numbers of water within the region of interest; CS is the contrast scale defined as CS = (?m – ?w)/(CTm – CTw), where ?m and ?w are the linear attenuation coefficients for the subject material and water respectively and CTm and CTw are the measured CT numbers of the subject material3. Uploading the acquired images into image-owl.
This system does not need any software to be installed with the current imaging system in princess marina hospital. The acquired images were uploaded into image owl which immediately gave a complete report both graphically (graphs) and numerically (tables with figures) on the uploaded images.
4. Interpreting results from image owl, drawing conclusions and relevant recommendations were made accordingly.

5. Noting down the values of CTDI shown on the computer and water. CS will be determined using CT number linearity phantom.