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Tomographic Imaging

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Tomographic Imaging. SPECT PET Hybrid. Spect. Conventional, Planar Imaging. Tomographic Imaging. Series of Projection images. Camera head(s) rotate about patient 360 o for most scans 180 o for cardiac scans Continuous acquisition or Step & Shoot Projection images acquired - PowerPoint PPT Presentation
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Tomographic Imaging

Tomographic ImagingSPECTPETHybridSpect

Series of Projection images

Conventional, Planar ImagingTomographic ImagingCross-section of the response characteristics of an idealized gamma camera. Each collimator hole views the radioactivity within a cylinder perpendicular to the face of the gamma camera, called its line of response. Under idealized conditions (such as no attenuation or scatter) the signal recorded by the detector at that point reflects the sum of activity within the line of response. For a row of holes across the detector, the gamma camera generates a projection profile as shown. The projection profiles provide the data from which the image is reconstructed

Rotating the gamma camera around the object provides a set of one-dimensional projection profiles for a two-dimensional object, which are used to calculate the two-dimensional distribution of radioactivity in the object. ECT, emission computed tomography.

3Data AcquisitionCamera head(s) rotate about patient360o for most scans180o for cardiac scansContinuous acquisition or Step & ShootProjection images acquiredImages reconstructedFiltered BackprojectionDirect Fourier TransformIterativeMulti-slice imaging

A planar gamma camera image is a projection of a 3D radioactivity distribution onto a 2D plane. As a result, it is difficult to extract detailed information about an object of interest from the superposition of overlying or underlying radioactivity.The goal of image reconstruction methods in SPECT is to estimate accurately the 3D radioactivity distribution in vivo without the overlapping information. SPECT imaging systems are designed to acquire projection data (similar to those obtained with the conventional planarmethod) accurately and efficiently from multiple views around the patient. Image reconstruction methods are applied to the projection data to generate SPECT images of the 3D distribution of radioactivity.1996 RadioGraphics, 16, 173-183.According to the theory of CT, projection views acquired over only 180 of arc are required for correct reconstruction. In a perfect imaging system, projections opposite each other are essentially mirror images of each other. Thus, the opposing views are redundant, and only 1 is needed. However, the nuclear medicine gamma camera is not a perfect imaging system, therefore, opposing views are not the same. First, the resolution of the gamma camera degrades as the distance between the camera and object being imaged increases. Second, a certain percentage of Compton scatter is accepted as photopeak gamma rays, due to the finite energy resolution of the camera. Third, a certain fraction of gamma rays from an object is attenuated (absorbed) when they are emitted in an attenuating medium, such as a patient. This phenomenon varies according to the depth of attenuating medium between the object and the gamma camera. In clinical SPECT, opposing projection views will never be the same. Therefore, 360 of arc is required for accurate reconstruction in most SPECT studies. J. Nucl. Med. Technol. December 1, 2000 vol. 28 no. 4 233-244

4Cardiac Scan

Myocardial perfusion studies are acquired with 180o arc.

Projection images from opposite 180o have poor spatial resolution & contrast due to greater distance & attenuation.

Myocardial perfusion studies are acquired with 180o arc as shown above. The projection images from the opposite 180o have poor spatial resolution & contrast due to greater distance & attenuation.5Multi-Head SPECT Systems

A dual-headed gamma camera system (top).

Note that the camera heads can be placed in different orientations to provide 2 simultaneous views of an organ or the body (bottom).Typically 180 for whole body SPECT, 90 for cardiac imaging

Sensitivity 2 angular projections acquired simultaneously 2-fold total # countsORSame # of counts acquired in time

Orbit Shape

The SPECT imaging system rotates around the long axis of the patient, who is lying on the SPECT imaging table. The radius of rotation is adjusted so that the camera will not come into contact with either the patient's surface or the table. For circular orbits, this places the gamma camera head(s) far from the patient in the anterior and posterior projections. To avoid this problem, the gamma camera can determine an orbit, either automatically or with the aid of the technologist, which will bring the camera as close to the patient as possible at all angles, improving spatial resolution . Then, during the actual SPECT acquisition, the camera will move in and out radially as it rotates around the patient (hence, the name noncircular or body contouring orbit).

7Image Reconstruction

2-D intensity display of set of projection profiles (sinogram)

Each row in display corresponds to individual projection profile, sequentially displayed from top to bottom.

Point source of radioactivity traces out a sinusoidal path in the sinogramProjection views are acquired at evenly spaced angles around the long axis of the patient, resulting in images with rows and columns of equidistant sampled areas . Pixels represent summations of the voxels at an angle perpendicular to the camera face. In the computer, the 3-D volume of radioactivity representing a function of activity vs. 3-D position in space, is viewed as a stack of 2-D transverse slices of thickness equal to the z dimension of a voxel, as illustrated. In filtered backprojection reconstruction, each row of each projection image is viewed as a 1-D representation of the object's projection. The 1-D pixel profiles are modified mathematically, or filtered, and projected back across the two-dimensional slice at their respective angles (ergo backprojection). The 2-D profile backprojections are added together, forming the reconstructed 2-D transverse image.

8

Original, ideal frequency spectrum of a 1-D SPECT projection profile

(B) Frequency spectrum (MTF), of SPECT imaging system (i.e., gamma camera/collimator).

(C) White noise frequency spectra for 3 levels of Poisson statistical noise (1 = lowest level). (D) Original profile spectrum is multiplied by SPECT system MTF & 3 noise spectra are added. Spatial frequency where the object signal falls to level of noise increases as noise level decreases (i.e., as acquired counts increase).

Spatial frequencies in the case of SPECT, refer to the frequencies contained in the variation of counts corresponding to objects (organs, tumors) in the patient, for example small objects and sharp edges contain more high frequencies than do broad, flat objects. Spatial information is converted to frequency information by the mathematical process known as the Fourier transform, or FT.

The spatial frequencies of SPECT data are digitally sampled, as are the images themselves. What spatial frequencies are contained in SPECT images? The answer lies in the gamma camera system's (including collimator) ability to detect high frequencies (small objects), and how finely the data coming from the camera is sampled (number of pixels).

As discussed earlier, a three-dimensional distribution of radioactivity is reconstructed as a stack of two-dimensional transverse slices of finite thickness, through backprojection of a series of two-dimensional projections. The projection profiles are smoothed, or blurred, by the projection process. To reconstruct the original, unblurred three-dimensional distribution, the profiles must first be filtered by a function known as a linear ramp in the spatial frequency domain. The linear ramp can simply be thought of as an amplifier, with increasing amplification as the frequencies increase, boosting the ability to see small objects (higher frequency objects, a higher power telescope). The linear ramp is a necessary, compensatory filter, as it removes the blurring effect of the projection process. In clinical SPECT, however, 2 problems arise. First, due to its finite resolution, the gamma camera/collimator imaging system is a low-pass filter, reducing the amplitude of the projection profile's frequency spectrum as the frequency increases. Thus, only a smoothed version of the original three-dimensional distribution at best may be reconstructed to begin with. Second, clinical nuclear medicine images tend to be photon deficient. The Poisson statistical noise inherent in all nuclear medicine images has approximately the same amplitude at all frequencies, and is known as white noise. The Poisson statistical noise is added to the already blurred profiles, and the final result is a blurred, noisy profile.

91/r Blurring

Computer-simulation phantom

B. Sinogram of simulated data for a scan of the phantom

C. Image for simple backprojection of data from 256 projection angles. 1/r blurring is apparent in the object, and edge details are lost.A,Computer-simulation phantom used for testing reconstruction algorithms. B, Sinogram of simulated data for a scan of the phantom. C, Image of simulation phantom for simple backprojection of data from 256 projection angles. 1/r blurring is apparent in the object, and edge details are lost.10Filter Kernels

Ideal Ramp FilterRamp Filters w/ Roll-OffRemoves 1/r blurring, sharpening image detailAmplifies high frequency noiseStatistical noise (random nature of decay & photon interactions) dominates high frequencies roll-off will smooth imageRamp filter in the spatial-frequency (k-space) domain. The filter selectively amplifies high-frequency components relative to low-frequency components. The filter removes the 1/r blurring present in simple backprojection and sharpens image detail, but it also amplifies high-frequency noise components in the image

Since statistical noise in image due to random nature of radioactive decay & photon interactions dominates the high frequency realm the ramp filter must be rolled off at higher frequencies to smooth the image.

11Iterative Reconstruction

More computationally intense than FBPRequires > 1 iteration / imageEach iteration 1 FBPAlgorithms often incorporate characteristics of imaging deviceCollimator & object scatterSystem geometryFinite detector resolutionSchematic illustration of the steps in iterative reconstruction. An initial image estimate is made and projections that would have been recorded from the initial estimate then are calculated by forward projection. The calculated forward projection profiles for the estimated image are compared to the profiles actually recorded from the object and the difference is used to modify the estimated image to provide a closer match. The process is repeated until the difference between the calculated profiles for successively estimated images and the actually observed profiles reaches some acceptably small level.12Iterative Example

Brain images generated for different numbers of iterations by an iterative reconstruction algorithm. Image resolution progressively improves as the number of iterations increases. In practice, the iterations are performed until an acceptable level of detail is achieved or until further iterations produce negligible improvement13Matrix Size

An important aspect of SPECT imaging is the selection of the matrix size in the computer for the projection views. Essentially, the computer divides up the gamma camera field of view (FOV) into square areas (pixels), and the 2 matrix sizes typically associated with SPECT imaging are 64 64 and 128 128, rows by columns. The choice of matrix size depends on several factors. First, the size of a pixel should, ideally, be less than 1/3 of the expected full-width at half-maximum (FWHM) resolution of the SPECT system, measured at the center of rotation for the isotope being imaged, including the effects of the collimator and the radius of rotation (i.e., distance of camera from patient). Higher SPECT resolution will always be achieved with the smaller pixel size of 128 x 128 matrices, however pixel SNR may be much poorer as the counts are divided up into 4 times the pixels of a 64 x 64 matrix image covering the gamma camera FOV. For the same acquisition, a 128 x 128 image only has 1/4 the counts per pixel.Ideally, for accurate reconstruction the number of angular views over 360 should be at least equal to the projection image matrix size (e.g., 64 views for a 64 64 matrix and 128 views for a 128 128 matrix). When the number of views is less than the minimum, streak artifacts may appear in the reconstructed slices.14Sampling Effects

Linear Sampling of Projections# of Angular Samplesr (Sampling Interval) FWHM/3Nviews FOV/(2r )LinearImages of a computer-simulation phantom reconstructed with progressively coarser sampling of the image profiles. Linear undersampling results both in loss of resolution and image artifacts

AngularEffect of the number of angular samples recorded on the reconstructed image of a computer-simulation phantom. Spoke-like streak artifacts are evident when an inadequate number of projections are used.

15Sampling Coverage

Effects of angular sampling range on images of a computer-simulation phantom. Images obtained by sampling over 45, 90, 135, and 180.

Sampling over an interval of less than 180 distorts the shape of the objects and creates artifacts

Effects of angular sampling range on images of a computer-simulation phantom. Images obtained by sampling over 45, 90, 135, and 180. Sampling over an interval of less than 180 distorts the shape of the objects and creates artifacts16Insufficient FOV

Effects of having some profiles that do not cover the entire object.

Left, Sinogram of computer-simulation phantom. Right, Reconstructed image

Effects of having some profiles that do not cover the entire object. Left, Sinogram of computer-simulation phantom. Right, Reconstructed image17Detector Failure

Effects of missing projection elements on reconstructed image. Left, Sinogram of computer-simulation phantom. Right, Reconstructed image.

Effects of missing projection elements on reconstructed image. Left, Sinogram of computer-simulation phantom. Right, Reconstructed image. This simulation would apply to a SPECT image reconstructed from profiles acquired over a 180-degree sampling range with a single-headed camera, with one region of the detector dead.

18SNR ComparisonPlanar ImagingTomographic ImagingStronger requirements on counting statistics for tomographic imaging as compared with planar imaging to achieve same level of SNRCNR ComparisonPlanar ImagingTomographic ImagingFor the same level of object contrast & total # of image counts (in absence of distance & attenuation effects), no intrinsic difference in CNR between planar & tomographic imagingTomographic Imaging AdvantagesDetecting low-contrast lesionsAbility to remove confusing overlying structures that interfere w/ lesion detectabilitye.g. ribs overlying lesion in lungsLesion shape & borders also become clearerDoes not improve detectability of lesions by CNRMore accurate determination of radioactivity concentrations in particular tissue volumePlanar Vs. SPECT

Thoracic phantom images

Contrast & visibility when overlying activity removed in SPECTPlanar (upper left) and single-photon emission computed tomographic (SPECT) (center) images of a thoracic phantom. Note the improved contrast and visibility of the voids in the cardiac portion of the phantom when overlying and underlying activity are removed in the SPECT images. 22SPECT ChallengesActual LOR resembles diverging cone rather than cylinder

Signal recorded not exactly proportional to total activity w/in LORSignal from activity closer to detector more heavily weighted than deeper lying activity due to attenuation of overlying tissueActivity outside LOR contributes to signalCrosstalk due to scattered radiationSeptal penetration through collimator

Most of discrepancies from ideal vary w/ -ray energyLead to artifacts & can seriously degrade image qualityDivergence of Response Profile

Volumes of tissue viewed by a collimator hole at 2 different angles separated by 180.

Differences in the volumes viewed results in different projections from the 2 viewing anglesVolumes of tissue viewed by a collimator hole at two different angles separated by 180 degrees. Differences in the volumes viewed results in different projections from the two viewing angles24Attenuation Effects

Attenuation leads to further differences in these two projections, emphasizing activity that is close to the gamma camera compared with activity further away that has to penetrate more tissue to reach the gamma camera.

Values are shown for the attenuation of the 140-keV rays from 99mTc in waterConjugate Counting

Response profiles vs. source depth for single view projection of line source in air & H20 AIR: Degradation of spatial resolution w/ distance from collimatorH20: Degradation due to distance & attenuation2 opposing view projections of line source in air & water arithmetically averagedAIR: No degradation w/ distanceH20: Only degradation due to attenuation2 opposing view projections of line source in air & water geometrically averagedAIR: No degradation w/ distanceH20: No degradation w/ distance

Line-spread functions versus distance in air (left) and in water (right) for high-resolution parallel-hole collimator on a gamma camera.

Measurements were made either with the tank empty (in air) or filled with water. Top, single detector only; middle, arithmetic mean of opposing detector profiles; bottom, geometric mean of opposed detector profiles. 26Attenuation & Photon Energy

The effect of attenuation on SPECT imaging. The intensity of photons emerging from a source of activity within an attenuating medium is reduced by e-x, where is the linear attenuation coefficient (cm-1) and x is the depth of the activity in the attenuating medium at a particular projection angle. This produces the cupping artifact in the reconstructed transverse slices (lower right).

In the brain or abdomen, which are dominated by soft tissue, is approximately constant. The attenuation phenomenon is much more complicated in the thorax, where varies throughout the volume (e.g., soft tissue, lungs, bone).Attenuation correction methods may be categorized as: constant , Chang methodAn attenuation map, based on patient boundary determination and an approximate or measured, constant , is generated and applied to the reconstructed transverse slices(b) variable , transmission source methodThe quantitative value of is determined by the use of transmission scanning, either with moving line sources, or fixed sources of varying geometries. From the transmission scan, a map is generated, the inverse of which provides the attenuation correction factor. The transmission sources are usually long-lived isotopes of dissimilar energy to both 201Tl and 99mTc photons (usually 153Gd, with a nominal 100-keV photon), and therefore a correction factor must be applied to convert the attenuation coefficient of the transmission source energy to that of 201Tl or 99mTc.27Attenuation Correction

If attenuation coefficient constant throughout tissue volume(reasonable assumption in brain, abdomen)Flood sourceLine sourceIf attenuation coefficient not constant throughout tissue volume(reasonable assumption in thorax, pelvis)Transmission scanChang MethodSPECT images of a 20-cm diameter cylinder containing a uniform concentration of 99mTc with and without attenuation correction (Chang method with narrow-beam attenuation coefficient of = 0.15cm1). Profiles are through the center of the images. The apparent overcorrection of attenuation is due to scattered events in the dataset.

Transmission ScanThe black arrows show the direction of rays emitted from collimated transmission source; the gray arrows show the direction of motion of moving line sources. A, Flood source. B, Collimated moving line source28Attenuation Map

Attenuation map of the thorax reconstructed from the reference and transmission scans obtained with a moving line transmission source

Reference Scan: 1st scan acquired w/ no object in FOV

Transmission Scan: 2nd scan acquired w/ object of interest in FOVScatter Correction

Dual energy windows used to simultaneously acquire SPECT & transmission scans

Patient equivalent phantom scanned to acquire scatter distribution in projectionsDual-energy windows used to simultaneously acquire SPECT emission (99mTc) and transmission (153Gd) data. Note the presence of downscatter from the 99mTc activity in the 153Gd window. The magnitude of the downscatter contamination depends on the relative amounts of 99mTc activity in the body and 153Gd activity in the transmission source, the amount of scattering material in the FOV, the details of how the transmission source is collimated, and the precise energy windows that are used

Diagrammatic sketch showing dual-energy windows superimposed on the spectral distribution of unscattered and scattered events for a patient-sized phantom filled with 99mTc30Partial Volume Effects

Each cylinder contains same concentration of radionuclide, but w/ diameterFor sources/volume > 2 x FWHM, image intensity reflects both the amount & concentration of activity w/in volume

For smaller objects that only partially fill voxel, total amount of activity still correct, but intensities of pixel no longer reflect concentration of activity

Spillover: when ROI has low tracer accumulation relative to surrounding tissues activity from these areas spills over to ROI

Results in contrast & under or over-estimation of tracer concentrationsSPECT Collimator Design

Parallel HoleFan BeamIn SPECT mainly use parallel hole collimators except for brain which uses fan beamFan BeamThey are designed for a rectangular camera head to image smaller organs like the brain and heart. When viewed from one direction, the holes are parallel. When viewed from the other direction, the holes converge. This arrangement allows the data from the patient to use the maximum surface of the crystal. When the Fanbeam is flipped over it is called a Single Pass Diverging Collimator used for whole body sweeps.

32Spatial Resolution

waterCo 57 line sourcesIn general the spatial resolution in SPECT is slightly worse than in planar imaging.Camera head farther from patientSpatial filtering used to reduce noise reduces resolutionShort time/view lower resolution collimator to obtain adequate numbers of counts

Tangential resolution at periphery > either center resolution or radial resolution at periphery.In general the spatial resolution in SPECT is slightly worse than in planar imaging.Camera head farther from patientSpatial filtering used to reduce noise reduces resolutionShort time/view lower resolution collimator to obtain adequate numbers of counts

PlanarRadioactivity in tissue in front of & behind tissue/organ of interest reduces contrastThis non-uniform pattern of radioactivity superimposes on activity distribution of tissue of interestStructural noiseSPECTImproved contrast/reduced structural noise by eliminating counts from activity on overlapping structuresIf iterative reconstruction implemented, can partially compensate for effects of scattering photons in patient and collimator effects-decreasing spatial resolution with distance from the camera andseptal penetrationWhen attenuation is measured with sealed source or CT data, can also partially correct for patient attenuation

33SPECT vs. PlanarPlanarRadioactivity in tissue in front of & behind tissue/organ of interest contrastNon-uniform pattern of radioactivity superimposes on activity distribution of tissue of interestStructural noiseSPECT contrast & structural noise by eliminating counts from activity on overlapping structuresIf iterative reconstruction implemented: partially compensate for effects of scattering photons in patientcollimator effects spatial resolution with distance from cameraseptal penetrationWhen attenuation is measured with sealed source or CT data, can also partially correct for patient attenuation

SPECT QCX & Y Magnification FactorsMulti-Energy Spatial RegistrationCenter of RotationMechanical COR must coincide w/ COR defined for each projectionIf detector sags/wobbles as it rotates, artifacts resultAdditional blurring or ring artifactsUniformityEven very small non-uniformities can lead to major artifacts unlike planar imaging)Rings or arcs in imagesFlood field uniformities 0, s not emitted at 180oAttenuation Correction511 keV annihilation 511 keV annihilation dd-xxProbability of both s escaping patient without interaction independent of where annihilation occurredAttenuation in PET differs from SPECT since both (annihilation) photons must escape patient to cause coincident event to be registered.

It is difficult to mathematically predict the specific appearance of PET images reconstructed without attenuation compensation. It is possible, however, to predict some common artifacts such as enhanced activity in pulmonary regions (that is, hot lungs) and negative tracer concentrations in mediastinal regions in uncorrected images. In regions of non-uniform density, such as the thorax, the lack of attenuation correction can mask the appearance of solid lesions with moderately elevated tracer uptake in the resultant images Seminars in Nuclear MedicineVolume 33, Issue 3, July 2003, Pages 166179

Probability that both photons will escape patient without interaction is the product of the individual probabilities:(e-x)(e-(d-x)) = e-d57Attenuation Correction: Transmission Measurement

rod source-Ge68-Ga68 + emitterAttenuation Methods

Ge68-Ga68 positron sourceCs-137 gamma ray source120 kVp x-ray source3 transmission methods used for measured attenuation correction for PET. In theseimages darker regions correspond to higher density (ie, bone).59With/Without Attenuation Correction

transmission (attenuation) image18FDG uptake with attenuation correctionwithout attenuation correctionImage Reconstruction

Filtered back projection or iterative methods can be usedFor 3D PET acquisition, iterative methods must be used

61Filtered Backprojection

Sinograms and ring gantry. Ring gantry is shown with 16 detector blocks. (B) In sinogram, events involving Block 12 are displayed along diagonal as shown. All blocks in coincidence with Block 12 (Blocks 17) are also displayed along diagonals, but slanting inopposite direction. LOR shown in as dashed line is represented in sinogram in (B) as intersection of dashed lines from Block 12 and Block 4, respectively. (C) If particular block in scanner is malfunctioning, it will lead to diagonal streak in sinogram as shown here. This fact is used in routine PET quality control to determine which blocks may need to be serviced.

J. Nucl. Med. Technol. June 1, 2002 vol. 30 no. 2 39-49

622D vs. 3D Acquisition

ABC

ABCA: Activity outside FOV REMOVEDB: Scattered Photon REMOVEDC: Valid Coincidence REMOVEDA: Activity outside FOV INCREASEDB: Scattered Photon INCREASEDC: Valid Coincidence ACCEPTED2DCoincidences are detected only within each ring/adjacent rings of detector elementsAxial septa (thin annular collimators typically Tungsten)Prevents most radiation emitted outside transaxial slice from reaching detector ring for that slice3DCoincidences detected between many/all detector ringsNo Septa usedNumber of true coincidences detected May permit smaller activities to be administeredDisadvantagesrandom coincidence fraction Dead time count losses Scatter coincidence fraction Number of interactions from outside FOV 63Coincidence Detection EfficiencyCoincidence detection efficiencyPosition along axis of PET2D3D2DEfficiency nearly constant along axial length of detector rings

3DEfficiency linearly from ends of rings to center 3D whole body acquisitions accomplished by discontinuous motion, greater overlap of bed positions necessary

2DEfficiency in detecting coincidence photons nearly constant along axial length of detector rings

3DEfficiency in detecting coincidence photons increases linearly from the ends of the rings to the center 3D whole body acquisitions accomplished by discontinuous motion, greater overlap of bed positions necessary64Time-Of-Flight

TOF

Ability of PET scanner to accurately measure time between 2 -interactions from 1 annihilation is defined as TOF capability.

Ability of PET scanner to accurately measure time between 2 -interactions from 1 annihilation is defined as TOF capability. If no TOF information is available (time resolution . 1,500 ps), probability that coincidence occurred along 1 LOR is basically same (A). However, if time resolution can be increased (,600800 ps), location of annihilation can be narrowed to several centimeters (B and C). In ideal PET system, where time resolution would be good (e.g., 15 ps), position of annihilation could be determined within several millimeters and would no longer require image reconstruction (D).

J Nucl Med June 2008 vol. 49 no. Suppl 2 5S-23S

66TOF Images

TOF PET images acquired with Phillips Gemini TF PET scanner. One can see that obese patient (119 kg; BMI, 46.5) particularly benefits from informationgained because of time resolution of 600 ps.

J Nucl Med June 2008 vol. 49 no. Suppl 2 5S-23S67PET QCTestDescriptionFrequencyUniformityUniform Scan of positron emitting sourceDailyTomographic UniformityScan of uniform cylindrical sourcePeriodicallyNormalizationMeasure efficiencies of all detector LOR & update stored normalization factorsQuarterly or if any uniformity test reveal non uniformityAbsolute ActivityIf quantitative measurements are to be usedQuarterlySystems Test:Spatial resolutionStatistical noiseCount rate performanceSensitivityImage qualityAnnually68Quantitative ImagingDesire: Pixel value to # of nuclear transformations Physiological Model developed in 1970sRate of local tissue glucose utilization calculated from amount of FDG that accumulates in tissueSUV: Standardized Uptake Valueg/cm3Attempts to normalize for:Administered activityRadioactive decayBody massFactors Limiting AccuracyAssayed activity accuracyExtravasation of activity during administrationAccuracy of attenuation correctionCorrect recording of elapsed timeAccuracy of patient body massPhysiological stateBody compositionSize of lesionMotionROI selectionPET ArtifactsAttenuation Correction Motion Stray Magnetic Fields Module Loss, Block Loss or Mis-calibration Coincidence Timing

Hybrid ModalitiesPET/CTPET/MRISPECT/CT

SPECT/CT

CT data can be used to correct for tissue attenuation in the SPECT scans on a slice-by-slice basis.

The integration of SPECT and CT systems into a single imaging unit sharing a common imaging table provides a significant advance in technology because this combination permits the acquisition of SPECT and CT data sequentially in a single patient study with the patient in an ideally fixed position. Thus, the 2 datasets can be acquired in a registered format by appropriate calibrations, permitting the acquisition of corresponding slices from the 2 modalities. The CT data can then be used to correct for tissue attenuation in the SPECT scans on a slice-by-slice basis. Because the CT data are acquired in a higher-resolution matrix than the SPECT data, it is necessary to decrease the resolution of the CT data to match that of SPECT.

J. Nucl. Med. Technol. March 2008 vol. 36 no. 1 1-10

73Attenuation Correction

Uncorrected SPECT scanAttenuation correction factorsAttenuation corrected SPECTArray of attenuation correction factors (B) can be determined from attenuation coefficient measurements determined from CT scan and used to correct emission counts from uncorrected SPECT scan (A) to provide final attenuation-corrected SPECT scan (C).

J. Nucl. Med. Technol. March 2008 vol. 36 no. 1 1-10

74Bilinear Model of CT attenuation Correction

CT has effective energy (mean) of approximately 70 keV, taking into account beam hardening. Because attenuation effects vary with energy, it is necessary to convert the attenuation data acquired with CT to match the energy of the radionuclide used in the SPECT acquisitions. For example, to convert the attenuation data measured at an effective energy of 70 keV to 140 keV for 99mTc, we typically accomplish this by using a bilinear model

J. Nucl. Med. Technol. March 2008 vol. 36 no. 1 1-10

75PET/CT & SPECT/CT AdvantagesSuperior Attenuation CorrectionHigh photon flux reduces statistical noiseImaging time reducedPost injection CT scans can be madeEliminates need for (consumable) transmission sourceAnatomic CT images fused with functional SPECT scanFunctional anatomic maps There are numerous advantages in the use of CT data for attenuation correction of emission data. First, the CT scan provides a high photon flux that significantly reduces the statistical noise associated with the correction in comparison to other techniques (i.e., radionuclides used as transmission sources). Also, because of the fast acquisition speed of CT scanners, the total imaging time is significantly reduced by using this technology. Another advantage related to the high photon flux of CT scanners is that attenuation measurements can be made in the presence of radionuclide distributions with negligible contributions from photons emitted by the radionuclides (i.e., postinjection CT measurements can be performed). The use of CT also eliminates the need for additional hardware and transmission sources that often must be replaced on a routine basis. And of course the anatomic images acquired with CT can be fused with the emission images to provide functional anatomic maps for accurate localization of radiopharmaceutical uptake. J. Nucl. Med. Technol. March 2008 vol. 36 no. 1 1-10

76PET/CT Artifacts

Respiration

A noiseless simulation of the effect of respiration on CT-based attenuation correction as compared with a standard PET transmission scan. In the PET emission data a 1.5-cm diameter lesion (4:1 contrast) has been simulated. The appearance of the lesion is blurred axially due to respiration during the PET emission scan. The CT scan was acquired during maximum inspiration which has a spatial mismatch with the respiratory averaged PET emission and transmission scans. The result of using the CT-based attenuation correction is an apparent axial shift of the top of the liver dome.

Seminars in Nuclear Medicine, Vol XXXIII, No 3 (July), 2003: pp 166-17978Contrast Agent

Effect of contrast agent accumulation. CT image showing regions of highly enhanced CT values due to a focal accumulation oral contrast in stomach (arrow). (b) PET image without attenuation correction. (c) PET image with CT-based attenuation correction.

Seminars in Nuclear Medicine, Vol XXXIII, No 3 (July), 2003: pp 166-179

79Truncation

The problem of truncated CT attenuation information for clinical PET/CT studies. The field of view for the CT and PET scanners are 45 and 60 cm, respectively. (a) Transaxial sections through CT whole-body image volume at the level of the lower liver. Transaxial sections through PET image before (b) and after (c) CT-based attenuation correction.Dashed vertical lines indicate the CT field of view.

Seminars in Nuclear Medicine, Vol XXXIII, No 3 (July), 2003: pp 166-179

80PET/MRI

PET/CT vs. PET/MRI

Images of 2 BALB/c mice bearing CT26 tumor. In combination PET/CT (first row), region of low tracer uptake in tumor (blue arrow) cannot be explained by CT images, because entire tumor appears as homogeneous tissue. In sharp contrast, T2-weighted MR images obtained by simultaneous PET/MRI reveal hyperintense area corresponding to area of low tracer uptake (red arrow). This is indication of tumor necrosis. From these images, one domain in which CT excels is also visible: inside lung, MR image depicts nearly no structures because of lack of signal-generating protons; in contrast, CT images reveal details of lung bronchia.

J Nucl Med June 2008 vol. 49 no. Suppl 2 5S-23S

82DoseModalityEffective Dose (mSv)PET/CT-10 mCi DosePET7CT diagnostic16CT nondiagnostic4July 2008 Radiology, 248, 254-263. A PET/CT test has two components: a PET scan and a CT, which are done together. The radiation exposure from CT has a very wide range depending on the type of the test, the area of the body scanned and the purpose of the test. In its simplest form, a CT scan is used only for the localization of abnormalities seen on a PET scan (non-diagnostic scan). The radiation dose from such a scan can be low (e.g. an effective dose of about 7 mSv for a whole body study). However, the effective dose from a high resolution diagnostic scan can be quite high (up to 30 mSv for a whole body CT scan). The effective dose from a PET scan is modest and depends on the activity of the injected FDG (18F-Fluoro deoxyglucose) and is typically 8 mSv for adults using 400 MBq and is the same whether a part of the body or the whole body is imaged. Major reductions in radiation doses from PET/CT scans can be achieved by modifying the acquisition parameters for CT.IAEA

83Effective Doses

84END85

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