[Medical Radiology] Radiation Oncology || Medical Imaging Modalities in Radiotherapy

Medical Imaging Modalities in Radiotherapy 625

C O N T E N T S

45.1 Computed Tomography 62545.1.1 Imaging Principles and Information Content 625 Image Acquisition and Reconstruction 625 Acquisition Parameters 626 Information Content 62645.1.2 CT Simulators 627 Specifi cations for Hardware and Software Components 627 CT Scanning and Image Interpretation Principles 62745.1.3 4D CT Imaging 628 Description and Clinical Applications 628 Acquisition and Reconstruction 62845.1.4 Linac Integrated CT and Cone-Beam CT 62945.1.5 Quality Assurance 630

45.2 Positron Emission Tomography–Computed Tomography (PET/CT) 63045.2.1 PET Principles and Information Content 630 PET Physical Basis and Data Acquisition 630 PET Data Corrections and Image Reconstruction 630 PET Image Content and Quantitation 63245.2.2 PET Performance Parameters 632 Spatial Resolution 632 Sensitivity 632 Noise-Equivalent Count Rate (NECR) 63245.2.3 PET/CT Integration 632 System Description 632 Clinical Advantages Over Independent PET and CT Scans 633 PET/CT Radiotherapy Simulator 63345.2.4 FDG Imaging 634 Underlying Mechanism 634 Imaging Protocol 634 Image Interpretation Principles 63445.2.5 4D PET 634 Description and Clinical Application 634 Acquisition and Reconstruction 63445.2.6 Quality Assurance and Radiation Safety 635

45.3 Magnetic Resonance Imaging 63545.3.1 Physical Basis and Instrumentation 63545.3.2 MRI Radiotherapy Applications 635

45.4 Outlook 63645.4.1 CT and CBCT 63645.4.2 PET/CT 63645.4.3 MRI 636

References 636

Medical Imaging Modalities in Radiotherapy 45Dimitre Hristov and Lei Xing

D. Hristov, PhD; L. Xing, PhDDepartment of Radiation Oncology, Stanford University School of Medicine, 875 Blake Wilbur Drive, Stanford, CA 94305-5847, USA

Introduction and Objectives

Therapeutic ratio improvement by exploiting the highly conformal distributions that are enabled by current image-guided radiation delivery systems depends on the accurate identifi cation and local-ization of both treatment targets and healthy structures. While imaging is indispensable in accomplishing this task, the refl ection of disease-related morphological and functional features in 3D and 4D medical image datasets is determined by: physical contrast generation mechanisms; imaging system design, performance, and capabilities; and clinical acquisition protocols. The intent of this chapter is to provide an overview of the above factors for the most widely used medical imaging modalities in radiotherapy. Thus this chapter examines, to varying degrees of detail:

Computed tomography (CT) �Positron emission tomography-computed tomography (PET/CT) �Magnetic resonance imaging (MRI) �

45.1 Computed Tomography

45.1.1 Imaging Principles and Information Content

Image Acquisition and Reconstruction

A modern multi-slice CT scanner and its major �components are illustrated in Figure 45.1. A gan-try-mounted fan pair comprising an X-ray source and detector array (Fig. 45.1) rotates continuously around a subject at between one and three rev-olutions per second as the subject is advanced through the scanner bore in continuous (spiral or helical scanning) or incremental (axial scanning)

LuBrady-BOOK.indb 625 13.08.2008 14:04:29

626 D. Hristov and L. Xing

fashion. Projection data containing the integral X-ray attenuation along paths connecting the X-ray source and the individual detectors within the array is measured (Fig. 45.2). These projec-tions are fi ltered with various convolution kernels depending on the selected clinical reconstruction protocol and then backprojected to reconstruct a CT image or slice (Oppelt 2006). Current multi-slice CT scanners use many (16– �320) detector rows allowing a number of slices to be measured simultaneously. This leads to re-duction in scan time and better utilization of the X-ray source output. However, depending on the clinical protocol only a subset of the detector rows may be used (Oppelt 2006).

Acquisition Parameters

Scan type: spiral (or helical) and axial. �X-ray tube high voltage measured in kVs. It con- �trols the X-ray source radiation exposure rate (for fi xed tube current) as well as the spectrum of the X-ray beam. The latter infl uences the radio-graphic contrast of the CT images. Typical high voltage values range between 80 and 140 kV. X-ray tube current measured in mA. It controls �the radiation exposure rate and thus the signal-to-noise ratio in the CT images. The latter di-rectly affect the detection of low-contrast objects (Sprawls 1992). While CT systems allow a large range from ~10 mA to ~500 mA, for given scan parameters the maximum value of the X-ray tube current is most often limited automatically by the fi nite X-ray tube heat load capacity. The product of the X-ray tube current and the scan exposure time (measured in mAs) is the major dose deter-mining parameter. Typical effective dose values range between 1–15 mSv depending on the acqui-sition technique and the body site (McCollough et al. 2008a).Pitch defi ned as the ratio of table travel per rota- �tion in millimeters divided by the beam collima-tion. Typical pitch values range between 0.1 and 2. While small pitch values improve the image resolution along the scan axis, larger pitch values result in smaller imaging doses.

Information Content

A CT image is a discrete 2D matrix representation �of the spatial distribution of the X-ray attenu-ation coeffi cient within a scanned object. For a

Fig. 45.1. A modern multi-slice CT scanner and its major component

Fig. 45.2. The physical principles of CT acquisition. The CT detector array registers projections comprising the subject integral attenuation along rays between individual detectors and the X-ray source



particular matrix element or pixel (i, j), the corre-sponding CT number CT#(i, j) measures the rela-tive value of the average attenuation coeffi cient μ( , )x yi j in Hounsfi eld units (HU), thus

CT i jx yi j water

water

#( , )( , )

=−⎛

⎝⎜⎞⎠⎟

1000μ μ

μ

where � μwater is adjusted so as to give water a pixel value of zero independent of the X-ray spectrum. A normal CT scale ranges from –1024 HU to 3071 HU. Extended CT scales are available as an off-line post-processing option for patients with metallic implants (Coolens and Childs 2003).The CT value of human tissues depends on the kV �setting for the CT scan. Representative CT values for some human tissues are given in Figure 45.3.A typical patient’s 3D CT data set comprises more �than 100 2D images (or slices), each of which con-tains 512×512 pixels. With 16 bits per pixel, the size of such dataset exceeds 50 megabytes. CT imaging is characterized by high spatial integ- �rity, excellent reproducibility, high cross sectional spatial resolution (<1 mm) and low contrast sensi-tivity (<0.5%) for typical imaging doses.

45.1.2 CT Simulators

Specifi cations for Hardware and Software Components

A CT simulator is a CT system with additional �hardware and software components to enable ra-diation therapy simulation and planning based on CT-generated model of the patient anatomy.A fl at table top is necessary to reproduce the treat- �ment position intended for radiation delivery.Large bore size (~80 cm) is required to provide �maximum fl exibility for patient setup. For in-stance, such a bore facilitates the CT simulation of breast patients on inclined breast boards.Extended fi eld-of-view (FOV) reconstruction is nec- �essary to capture the anatomy of large patients and provide skin contours for evaluation of the deliv-ered dose. Since extended FOV reconstruction re-quires projection extrapolation beyond the acquisi-tion FOV (~50 cm), the fi delity of the CT numbers in the periphery of the FOV needs to be examined.Localizing (movable) lasers are to be included �for facilitating the radiation therapist with the patient setup and skin marks.Large X-ray tube heat capacity exceeding 6 Mega �Heat Units (MHU) is necessary to enable four-dimensional respiratory correlated CT scans (4D CT) with clinically useful scan range as well thin slice CT scans.Large detector array (>16 rows) is necessary to �allow large scan coverage at high resolution. This is benefi cial for single breath-hold imaging as well as 4D imaging.Extended HU scale is benefi cial in scanning pa- �tients with metal implants.Respiratory gating option is to be included for �retrospective 4D CT or prospective gated CT ac-quisition.

CT Scanning and Image Interpretation Principles

Scan limits need to encompass all relevant anatomy �for accurate calculation of dose and dose derived indices such as dose-volume histograms (DVHs).Thin slice imaging is important for treatment sites �such as head and neck (H&N), prostate, and lung since the CT volume or digitally reconstructed radiographs (DRRs) derived from the CT are sub-sequently used as reference data in image-guided radiation delivery.Fig. 45.3. Typical CT values for some human tissues



Extended HU scale with proper HU-to-electron �density calibration can reduce dose calculation uncertainty for patients with metal implants. However, an extended scale does not necessarily eliminate image artifacts.Clinical scanning protocols need to be created �and used consistently since HU, depending on the kV settings and reconstruction algorithms, are converted to electron density for accurate dose calculation.Visualization and interpretation of CT images de- �pend on the acquisition and the reconstruction parameters as well as the display setting used for viewing the data. The apparent size of a structure depends on the window-level settings. These need to be standardized and followed consistently.

45.1.3 4D CT Imaging

Description and Clinical Applications

4D CT imaging refers to CT imaging techniques �that allow the acquisition of respiratory-correlated scans (Vedam et al. 2003; Rietzel et al. 2005b; Pan et al. 2004; Keall et al. 2004; Li et al. 2005). Such scans can be acquired by prospective gating whereby imaging is performed only during a pre-determined respiratory state or by retrospective re-spiratory correlated sorting of CT images that cap-ture several states of the breathing cycle. The latter approach is referred to as retrospective 4D CT.A retrospective 4D CT allows: (i) mitigation of �image artifacts caused by respiratory-correlated internal anatomy motion; (ii) evaluation of the pattern and magnitude of the internal motion; and (iii) design of strategies for its management in the course of radiotherapy. These strategies include generation of patient-specifi c margins for target volumes ( Underberg et al. 2004, 2005; Rietzel et al. 2005a), breath-hold radiation deliv-ery (Wong et al. 1999), or gated delivery (Vedam et al. 2001, Kubo and Hill 1996, Ohara et al. 1989, Wink et al. 2008). 4D CT forms the basis for respiratory motion management in radiation therapy (Keall et al. 2006; Wink et al. 2008).

Acquisition and Reconstruction

The retrospective 4D CT scanning process in- �cludes three relatively independent steps: record-

ing of respiratory signal(s), acquisition of time-dependent CT projection data, and construction of a 4D image from these data. 4D CT patient setup proceeds along the same �lines as a standard 3D CT exam. The patient is immobilized on the scanner bed and aligned us-ing room and scanner lasers. Sagittal and coronal scout images are used to verify patient position-ing, and the setup is adjusted as necessary. At this stage of the setup, the 4D procedure begins to diverge from the 3D exam. Respiratory signal is recorded by tracking a sur- �rogate of respiration-related organ and tumor motion, such as chest expansion monitored by pneumatic bellows (Kleshneva et al. 2006) or displacements of a refl ecting external marker placed on the abdomen and tracked with a cam-era (Pan et al. 2004) (Fig. 45.4).Once a suffi ciently regular breathing pattern is �established, time-stamped CT data is acquired in either over-sampled helical (pitch ~0.1) or “cine” mode. The latter is a step-and-shoot technique, whereby the gantry completes several rotations at each bed position in order to acquire data over the full respiratory cycle. With either mode, several CT slices are generated that capture the anatomy over the full respiratory cycle at each axial loca-tion. Because several respiratory points are sam-pled at each bed position, a 4D CT scan can take several times as long as a corresponding 3D CT. A 4D CT scan typically results in 1500–3000 CT slices for a 20- to 40-cm axial FOV.Upon scan completion, phase or amplitude at each �point of the respiratory trace is calculated. In the

Fig. 45.4. 4D CT acquisition with an infrared camera track-ing system consisting of an infrared source, CCD camera, and a refl ecting block. The block is attached to the patient’s abdomen, typically just inferior to the xiphoid process, and the motion of the block is captured by the camera



case of phase calculation, the location of the peaks at end-inspiration is determined, and percentages to inter-peak points are assigned by a linear in-terpolation of the peak-to-peak distance. For ex-ample, under this scheme, end-inspiration occurs at 0%, while end-expiration typically appears near 50%–60%. The peak-to-peak distance can vary be-tween respiratory cycles, as can the position of end-expiration with respect to end-inspiration.The respiratory and scan data are combined by �sorting the over-sampled time-stamped CT slices according to their phase. Thus different CT series labeled in accordance with the respiratory state are generated (Fig. 45.5). These form the basis for 4D treatment planning.4D CT effective doses are about a factor of 5–10 �larger than those for standard thorax exams (Li et al. 2005).

45.1.4 Linac Integrated CT and Cone-Beam CT

Cone-beam CT (CBCT) is an imaging modality �which employs a large area (cone) X-ray beam and a fl at panel detector technology for the reconstruc-tion of 3D dataset from a number of 2D projections acquired from a subject (Xing et al. 2006).While analogous to CT, CBCT differs in two main �aspects: detector technology and collimation of the imaging beam impinging on the subject. For comparable imaging doses, these factors result

in CBCT image quality somewhat inferior to that of CT. Linac-integrated kV CBCT refers to a combination �a kV range X-ray source and a fl at-panel detector mounted on the drum of a medical accelerator with the kV imaging axis orthogonal to that of MV therapy beam (Jaffray and Siewerdsen 2000). Linac-integrated megavoltage CBCT refers to an �imaging mode of the linac delivery system that employs the megavoltage treatment beam as an imaging source in combination with a fl at-panel detector mounted opposing to the treatment source (Pouliot et al. 2005). The physics of radi-ation-matter interaction at megavoltage energies affects the MV CBCT image quality. Linac-integrated CT refers to a dedicated deliv- �ery system that integrates a 6 MV MV treatment source and an array of CT detectors on a CT gan-try (Meeks et al. 2005; Langen et al. 2005a,b; Kupelian et al. 2005).3D CBCT (or CT) images are used for on-line verifi - �cation and correction of patient setup. The images are registered with the planning CT data through the use of either manual or automated 3D image registration software that calculates shifts in x-, y- and z-directions (depending on the manufacturer, rotations can also be included). The movements determined during the registration represent the required setup corrections that are applied by dis-placing the treatment couch (White et al. 2007; Moseley et al. 2007; Letourneau et al. 2005; Gayou and Miften 2007; Xing et al. 2006).

Fig. 45.5. 4D CT reconstruc-tion by retrospective sorting. At each table position time-stamped images belonging to a particular respiratory state are selected and combined in a separate 3D dataset cor-responding to the respiratory state. (In this particular ex-ample the respiratory state is labeled by a respiratory phase. The 50% phase and the 90% phase are reconstructed)



45.1.5 Quality Assurance

CT scanner quality assurance programs need to �target the minimum standards for dose and im-age quality established by the American College of Radiology (ACR) Computed Tomography Ac-creditation Program. Quality assurance guidelines for CT simulators �covering radiation and patient safety, electrome-chanical components, as well as imaging perfor-mance are provided in an American Association of Physicists in Medicine (AAPM) report by Task Group 66 (Mutic et al. 2003).CBCT quality assurance, methods, and guide- �lines are described in (Letourneau et al. 2007; Mao et al. 2008; Yoo et al. 2006).

45.2 Positron Emission Tomography– Computed Tomography (PET/CT)

45.2.1 PET Principles and Information Content

PET Physical Basis and Data Acquisition

A radioactive isotope (Table 45.1) typically con- �jugated to some molecule of biological interest, decays via positron emission (Fig. 45.6a). The emitted positron travels some distance based on its energy (<1 mm–4 mm) before it encounters an electron and annihilates. The annihilation produces two 511 keV photons at approximately 180 , minus some small angle (~0.5 ) due to the energy of the positron (Fig. 45.6b). The 511 keV photons are stopped by scintilla- �tion detectors coupled to photomultiplier tubes (Fig. 45.6b). The light signal produced in the

scintillation detectors decays with characteristic times ranging from 40–300 ns depending on the detector scintillation material. After energy discrimination (between ~370 keV �and ~650 keV, depending on the scintillating material), a detected photon (single) that quali-fi es as an annihilation -ray is time stamped. A coincidence event is registered by identifying detectors that count a single energy-qualifying photon within a “coincidence window” of 10–20 ns. Thus each coincidence event is assigned to a particular line of response (LOR) which is the volume spanned by a pair of coincidence detec-tors (Fig. 45.6d).

PET Data Corrections and Image Reconstruction

The fi nite processing time (dead time) associ- �ated with the detection of a -ray results in loss of coincidence events. Known relationships be-tween measured and true events are used to es-timate true count rates from the detected ones (Townsend 2004, 2006; Steven and Badawi 2004; Cherry et al. 2003).Accidental (“randoms”) coincidences result from �separate electron-positron annihilation events when the photons originating in such events are registered in the coincidence window (Fig. 45.7). Since these events are temporarily uncorrelated, their number is estimated in a delayed coinci-dence window in which there are no true coinci-dence events. The random coincidence correction is a subtraction of the delayed-window counts from the coincidence window counts corrects for each LOR (Townsend 2004, 2006; Steven and Badawi 2004; Cherry et al. 2003).Non-uniform response of the detector elements �results in different LOR count rates for the same activity. For each LOR, the raw count rates are corrected by normalization factors estimated by scanning a positron-emitting source that exposes the detector pairs to uniform photon fl ux.When an annihilation � -ray is scattered, a coin-cidence event is registered in a misplaced LOR (Fig. 45.7) if the -ray falls within the energy dis-crimination window. Such scatter events result in image quality deterioration manifested in diffuse background counts. Scatter correction is build via scatter models into the reconstruction process (Steven and Badawi 2004).

Table 45.1. Some radioisotopes for PET imaging

Isotope Half life(h)

Decay modeabundance(%)

Maximum + energy

(keV)

Mean + energy

(keV)11C 0.34 + (99.8%) 960.2 385.613N 0.17 + (99.8%) 1198.4 491.815O 0.03 + (99.9%) 1731.9 735.318F 1.83 + (96.7%) 633.5 249.8



Annihilation events in a particular spatial location �result in different coincidence counts along differ-ent LORs because the photon fl ux attenuation varies from LOR to LOR. The attenuation factor is given by

exp( ( , ) )−∫ μ x y dlL

0 which measures the total atten-

uation at 511 keV along a LOR of length L . The attenuation corrections are derived either from transmission scans employing a single radioac-tive source (Cs-137 or Ge-68) or from attenuation maps reconstructed from the 3D dataset provided by the CT scanner in a PET/CT system (Steven and Badawi 2004).PET image reconstruction is based on iterative �algorithms that refi ne estimates of the activ-ity distribution by optimizing a target function which incorporates models of the data acquisi-tion process, statistical noise models and prior constraints such as the non-negativity of the count values (Boellaard et al. 2001; Hudson and Larkin 1994; Riddell et al. 2001; Cherry et al. 1992; Yao et al. 2000).

Fig. 45.6a–f. The principles of PET imaging shown schematically: (a) the decay of a neutron-defi cient, positron emitting iso-tope; (b) the detection in coincidence of the annihilation photons within a time window of 2 (10–20 ns); (c) the glucose analogue deoxyglucose labeled with the positron emitter 18F to form the radiopharmaceutical FDG; (d) the injection of the labeled phar-maceutical and the detection of a pair of annihilation photons in coincidence by a multi-ring PET camera. A line of response is shown; (e) the collection of the positron annihilation events into sinograms wherein each element of the sinogram contains the number of annihilations in a specifi c projection direction; and (f) a coronal section of the fi nal, reconstructed whole-body image mapping the utilization of glucose throughout the patient. [Reproduced from Townsend (2006) with permission]

a b c

d e f

Fig. 45.7. True coincidence (solid black), random coincidence (dotted black), and scatter coincidence (dotted red) events as-signed to one and the same line of response (thick red line)



PET Image Content and Quantitation

Reconstructed PET images represent count rate �(counts per second, or cps) per voxel provided that PET raw data is corrected for deadtime, randoms, detector response, scatter, and attenuation.With the application of a measured system cali- �bration factor given in (Bq/cc)/(cps/voxel), the PET images yield activity concentration A(i,j,k) ([A(i, j, k)]=Bq/cc) for each voxel (i, j, k). The sys-tem calibration factor is determined by a PET scan of water-fi lled volume source with uniform activity concentration. Another clinically widely used representation of �the activity concentration is the standard uptake value (SUV) which, under specifi c conditions, can approximate the net rate of radiotracer fl ux into the tissue (Weber 2005). SUV is defi ned as:

SUV decay corrected activity/tissue volumeinjected activit

=yy/body mass

and measured in g/cc.

45.2.2 PET Performance Parameters

Spatial Resolution

Positrons are emitted over a spectrum of kinetic �energies with maximum energies ranging from 0.58 MeV to 1.73 MeV depending on the radio-isotope. The fi nite travel range of positrons prior to annihilation degrades the spatial resolution by ~0.1 mm for F-18 and 0.5 mm for O-15 (Levin and Hoffman 1999, 2000).Non-colinearity of the annihilation photons re- �sults in blurring that depends on the PET scanner bore diameter. This spatial resolution degrada-tion is about ~2 mm for an 80-cm bore scanner (Zanzonico 2004; Cherry et al. 2003).Finite detector size (~4–6×4–6×20–30 mm � 3) and depth-of-interaction effects directly affect spa-tial resolution in radially dependent manner (Cherry et al. 2003).The overall spatial resolution of last generation �commercial whole-body PET scanners as mea-sured by full width half maximum (FWHM) of the line spread function is in the range of 4 mm–8 mm (Teras et al. 2007; Surti et al. 2007; Bettinardi et al. 2004).

Sensitivity

System sensitivity is defi ned by the measured �count rate per unit activity. It depends on two parameters. The fi rst one is the scanner geometric effi ciency as determined by the fraction of emit-ted photons striking the detector. The second one is the detector quantum detection effi ciency as determined by the fraction of photons striking a detector that are stopped and counted by the detector. System sensitivity is a measure of the utilization of the injected activity for imaging.Representative value for last generation com- �mercial whole-body PET scanners range between 3–7 cps/kBq (Teras et al. 2007; Surti et al. 2007; Bettinardi et al. 2004).

Noise-Equivalent Count Rate (NECR)

The noise-equivalent count rate (NECR) �(NU2-2001, 2001) is defi ned as

NECR TT S R

=+ +

2

, where T, S, and R are True,

scatter and random count rates. NECR provides a metrics for comparing different �PET systems and acquisition modes by equating count rates to the count rate that would have re-sulted in the same signal-to-noise ratio in the data in the absence of randoms and scatter.

45.2.3 PET/CT Integration

System Description

A PET/CT system integrates a CT scanner and a �PET scanner via a common which transfers a sub-ject between the imaging bores of axially aligned CT and PET scanners by simple linear motion (Townsend and Cherry 2001; Townsend et al. 2003). The known couch positions during the CT and the PET scan form the basis for the hard-ware fusion between the PET and CT datasets (Fig. 45.8).Additional software integration incorporates CT �data in the calculation of attenuation maps and other radiological properties necessary for the attenuation and scatter correction of the PET raw data (Fig. 45.9d).



Clinical Advantages Over Independent PET and CT Scans

PET/CT is a hardware-based image-fusion technol- �ogy that virtually eliminates the uncertainty and inconvenience of the software fusion of separate PET and CT images, which are often acquired with the patients in different positions during the two exams. Thus the PET/CT improves the accuracy of the anatomical localization of the metabolic signal and therefore the physician confi dence in making diagnostic and treatment decisions.Incorporation of CT attenuation maps reduces �whole-body scan times by about 40% (Townsend 2004).

PET/CT is more convenient for the patient and �the physician in comparison to a combination of separate PET and CT exams.

PET/CT Radiotherapy Simulator

A PET/CT simulator is a PET/CT system with ad- �ditional hardware and software components to enable radiation therapy simulation and plan-ning based on CT-generated model of the patient anatomy and PET-generated models of biological processes.Hardware and software requirements with re- �spect to the CT subsystem are equivalent to these for a CT simulator.

Fig. 45.8. A hybrid PET/CT scanner. [Adapted from Townsend (2006) with permission]

Fig. 45.9a–f. A typi-cal imaging protocol for a combined PET/CT study that comprises (a) a topogram, or scout scan, for positioning; (b) a spiral CT scan; (c) a PET scan over the same axial range as the CT scan; (d) the generation of CT-based attenuation correction factors; (e) reconstruction of the attenuation-corrected PET emission data; and (f) display of the fi nal fused images. [Repro-duced from Townsend (2006) with permission]

a topogram b spiral CT

d attenuation correction

c PET acquisition e PET reconstruction f fused PET/CT

FOREAWOSEM

Fusion

PETCT

PETCT



A fl at table top is necessary to reproduce the in- �tended treatment position during radiation de-livery.Depending on the design, the extended range of �the patient support system (couch) can lead to couch fl ex between the PET and CT exams. This effect results in some uncertainty of the hardware fusion process. Localizing (movable) lasers are to be included �for facilitating the radiation therapist with the patient setup and skin marks.Additional facilities such as injection/patient �waiting room in the vicinity of the PET/CT room need to be available.

45.2.4 FDG Imaging

Underlying Mechanism

By far the most commonly used radiotracer for �diagnosis, staging, detection of recurrent disease and monitoring of cancer therapy is the fl uorine-labeled glucose analog [18F]2-fl uoro-2-deoxy-D-glucose (FDG, shown in Fig. 45.6c) (Fletcher et al. 2008; Krause et al. 2007).Once inside a cell, FDG is phosphorylated to FDG- �6-phosphate but it is not metabolized beyond this step because of the fl uorine substitution in the molecule (Reivich et al. 1979). Compared to normal tissues many tumors are �characterized by glucose avidity (Warburg 1956) which results in preferential uptake of FDG and therefore in elevated accumulation of the emitter activity imaged by PET.

Imaging Protocol

Patients are instructed to have no caloric intake �for at least 4 h before imaging.FDG activity of about 5 MBq per kilogram of body �weight is administered, with typical values rang-ing between 370 MBq and 740 MBq.During a tracer uptake phase lasting between 45– �120 min the patient sits in a quiet resting room. After the uptake phase, the patient is placed �on the PET/CT table in a comfortable position that allows the entire anatomy of interest to be captured in the CT reconstruction fi eld-of-view. This minimizes errors in CT-based PET attenua-tion correction (Nestle et al. 2006). A topogram

(scout)view is acquired to determine the extent of the PET/CT scan (Fig. 45.9a–f.a).The patient is instructed to hold their breath at end �tidal volume or breathe shallowly to minimize the mismatch between the PET and the CT data. A spiral CT dataset is acquired followed by a PET �scan (Fig. 45.9a–f.a,b). PET images are reconstructed with and without CT attenuation correction. The lat-ter set of images facilitate interpreting and resolving ambiguities resulting from CT attenuation correc-tions in the presence of CT contrast, CT truncation artifacts, and metal implants (Nestle et al. 2006).

Image Interpretation Principles

Increased blood glucose level prior to scanning �results in lower FDG uptake.Increased tracer uptake time results in increased �FDG uptake.Imaged maximum activity concentration is mark- �edly underestimated when signifi cantly different uptake occurs across spatial scales smaller than twice the PET resolution. Examples are small le-sions and large tumors with necrotic centers and thin viable rims.Visualization interpretation of lesion size de- �pends on the display window-level settings.

45.2.5 4D PET

Description and Clinical Application

4D PET imaging refers to PET imaging techniques �that allow the acquisition of respiratory-corre-lated scans (Nehmeh et al. 2002, 2003; Klein et al. 1998; Huesman et al. 1997, 1998; Li et al. 2006b; Thorndyke et al. 2006).The most common solutions gated PET acquisi- �tion and list mode acquisition with retrospective reconstruction (Nehmeh et al. 2002, 2003; Klein et al. 1998; Huesman et al. 1997, 1998).A 4D PET scan allows: (i) evaluation of the pat- �tern and magnitude of the internal motion, and (ii) better estimation of the maximum activity uptake because of the elimination of motion blurring.

Acquisition and Reconstruction

For gated PET acquisition mode, patient setup pro- �ceeds in the same manner as an un-gated PET one.



For systems with an optical respiratory trace mon- �itor, an infrared camera tracks a refl ecting block placed on the patient abdomen (Fig. 45.4.).An acquisition trigger is set by the user to occur �at some given point (e.g., end inspiration) in the respiratory cycle. When this point is detected by the optical tracking system, a trigger is sent to the scanner and data accumulation is initiated. In gated mode, the user selects both the width �of the acquisition window and the number of se-quential bins to be recorded within each respira-tory cycle. The bin width directly affects image quality, since the signal-to-noise ratio within an image asymptotically approaches the square root of the signal level. Multiple bin acquisition allows capture of the full respiratory cycle in several bins, offering the possibility of retrospectively sorting into two or more respiratory phases.Each time a trigger is received, data is directed �to the initial bin, and then to the remaining bins sequentially until the next trigger. This process continues for the duration of the scan.

45.2.6 Quality Assurance and Radiation Safety

The quality assurance program for the CT sub- �system of a PET/CT scanner needs to target the minimum standards for dose and image quality established by the ACR Computed Tomography Accreditation Program. Quality assurance guidelines covering radiation �and patient safety, electromechanical compo-nents, as well as imaging performance provided in an AAPM report by Task Group 66 (Mutic et al. 2003) for CT simulators are applicable to PET/CT simulators as well.The quality assurance program for the PET sub- �system of a PET/CT scanner needs to target the minimum standards established by the ACCR Ac-creditation of Nuclear Medicine and PET Imag-ing Departments Program with respect to perfor-mance testing and quality control (MacFarlane 2006).Compliance with radiation safety standards �needs to be observed including adequate shield-ing of PET/CT facilities. Guidelines for shield-ing of PET/CT facilities are discussed in an AAPM report by Task Group 108 (Madsen et al. 2006).

45.3 Magnetic Resonance Imaging

45.3.1 Physical Basis and Instrumentation

Most MRI techniques involve generation and ma- �nipulation of bulk magnetization of soft tissue water protons within a given voxel through use of radio-frequency (RF) radiation and magnetic fi elds (McRobbie 2007). Magnitude, relaxation, and resonance properties of this bulk magne-tization can be interrogated by MR to serve as contrast generating mechanism. RF signal gener-ated by temporally changing bulk magnetization is detected, digitized, and processed in order to reconstruct the MR images. The major components of an MRI system include: �(1) magnetically shielded main magnet with a static fi eld strength between 0.2–3 Tesla (T) in order to generate bulk net magnetization along the magnetic fi eld direction; (2) gradient-coils creating linear variations of the main magnetic fi eld with maximum gradient values between 10–50 mT/m in order to spatially encode the bulk magnetization; (3) a radiofrequency system with a transmitting coil to tip the bulk magnetization away from the static magnetic fi eld direction and receiving coils to detect the radiofrequency signal resulting from the temporal variations of the bulk magnetization (Oppelt 2006; McRobbie 2007).With a patient positioned within the main mag- �netic fi eld, a typical MR acquisition protocol in-volves a sequence of gradient coil and transmitter coil activations along with RF signal detection by the receiving coils. In addition to the magnitude, relaxation, and resonance properties of the bulk magnetization, timing and order of coil activa-tions and readout signifi cantly infl uence the MR signal (Westbrook et al. 2005).

45.3.2 MRI Radiotherapy Applications

Typical MR protocols include several sequences to �generate images that are weighted with respect to proton density and tissue (magnetization) spin-lattice (T1) and spin–spin (T2) relaxation times (Villeirs and De Meerleer 2007; Jenkinson et al. 2007).



MRI provides superior soft tissue discrimination, �especially for CNS structures (Jenkinson et al. 2007; Henson et al. 2005) and within the abdo-men and pelvis (Villeirs and De Meerleer 2007; Khoo and Joon 2006).Fast MRI can be used for imaging temporal �variations of the anatomy to design appropriate motion management strategies in the course of radiotherapy (Chan et al. 2008; Plathow et al. 2006; Kauczor and Plathow 2006).MRI is typically employed together with CT im- �ages with the help of image fusion software to delineate the extent of the malignancy. Incorpo-ration of MR for the delineation of target volumes requires proper accounting for MR system and object-induced distortions, as well as anatomy variations resulting from differences between the MR and CT imaging setups.

45.4 Outlook

45.4.1 CT and CBCT

CT scanners with very large numbers of detec- �tor rows (256 or more) are becoming available that may enable 4D CT imaging of adequate anatomical extents without external respira-tory signal. With proper imaging dose man-agement, true anatomy motion evaluation even in the presence of irregular breathing patterns may be possible.Dual X-ray source scanners are being commer- �cialized (Engel et al. 2008; Flohr et al. 2006; McCollough et al. 2008b; Yan et al. 2006). These may offer the opportunity for radiotherapy rel-evant tissue characterization by dual-energy volumetric imaging.4D CBCT is being developed to obtain phase re- �solved volumetric images that eliminate motion artifacts. This can improve image interpretation and localization accuracy prior to treatment de-livery (Dietrich et al. 2006; Harsolia et al. 2008; Li et al. 2006a,c, 2007, 2008; Li and Xing 2007; Lu et al. 2007; Purdie et al. 2006; Rit et al. 2005; Sonke et al. 2005, 2008).

45.4.2 PET/CT

A number of radiotracers are being investigated �as agents for imaging biological processes and microenvironment parameters of interest to ra-diotherapy: proliferation, hypoxia, apoptosis, and angiogenesis [Nimmagadda et al. (2008) and references therein]. Fast scintillation detectors and new reconstruc- �tion algorithms are being pursued to enable time-of-fl ight PET imaging that will potentially improve image quality and PET quantifi cation (Karp et al. 2008; Surti et al. 2006, 2007).

45.4.3 MRI

Dynamic contrast enhanced MRI (DCE-MRI) is �being actively explored as a predictor of tumor response to radiotherapy [Zahra et al. (2007) and references therein].Magnetic resonance spectroscopy (MRS) is a fo- �cus of investigations as a potential tool for treat-ment planning and treatment response evalua-tion in brain tumors and prostate cancer [Payne and Leach (2006) and references therein].Available open-bore (>70 cm) MR systems outfi t- �ted with MR compatible fl at table tops and laser systems may be adopted as dedicated MR simu-lators.

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