8 Anato-Molecular Imaging: Combining Structure and Function

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8 Anato-Molecular Imaging: Combining Structure and Function

David W Townsend and Thomas Beyer

Introduction

Historical Perspectives

Non-invasive technologies that image different aspects of disease should really be viewed, in almost all cases, as complementary rather than competing. When the sensitivity and specificity of one imaging technique for diagnosing or staging a specific disease is compared to that of another technique, it is usually to establish the superiority of one of the two techniques. In practice, however, such comparisons are of little real value because anatomical and functional imaging techniques have different physical specifications of spatial, tempo- ral and contrast resolution, and the images even reflect different aspects of the disease process (Fig. 8.1). CT and MRI are used primarily for imaging anatomical changes associated with an underlying pathology, whereas the molecular imaging techniques of PET and

SPECT capture functional or metabolic changes associ- ated with that pathology. Historically, CT has been the anatomical imaging modality of choice for the diagno- sis and staging of malignant disease and monitoring the effects of therapy. However, more recently, molecu- lar imaging with whole-body PET has begun assuming an increasingly important role in the detection and treatment of cancer [1].

Nevertheless, historically, functional and anatomical imaging modalities have developed somewhat inde- pendently, at least from the hardware perspective. For example, until recently, CT developed as a single-slice modality, while PET and SPECT have always essentially been volume imaging modalities even if, for technical reasons, acquisition and reconstruction has been limited to two-dimensional transverse planes. CT de- tectors integrate the incoming photon ﬂux into an output current whereas PET detectors count individual photons. Some similarity can be found in the data pro- cessing as the image reconstruction techniques are

179

Figure 88.1. Transverse CT (a), PET (b), and MRI (c) image of the abdomen illustrating different aspects of the anatomy and metabolism of the patient as imaged with the different diagnostic imaging techniques.

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based on common theoretical principles. The imple- mentation details, however, involve many important differences that set the various modalities apart.

Consequently, over the past twenty-ﬁve years, the de- velopment of anatomical and molecular imaging tech- niques has followed distinct, but parallel, paths, each supporting its own medical speciality of radiology and nuclear medicine.

Image Fusion: A Hardware Approach

Complementing Anatomy with Function

In clinical practice a PET study, if available, is gener- ally read in conjunction with the corresponding CT scan, acquired on a different scanner and usually on a different day. Adjacent viewing of anatomical and func- tional images, even without accurate alignment and su- perposition, can help considerably in the interpretation of the studies. Using the retrospective software-based approaches anatomical and molecular images can be aligned and read as combined, or fused, images.

This can be an advantageous procedure because identification of a change in function without knowing accurately where it is localized, or equivalently, knowl- edge that there is an anatomical change without understanding the nature of the underlying cause, compromises the clinical efficacy of both, the anatomi- cal and functional imaging. More importantly, since a functional change may precede an anatomical change early in the disease process, there may be no identifiable anatomical correlate of the molecular change, although of course a sufficient number of cells must first be affected to produce a macroscopic change that can be imaged with a PET scanner.

Software Approach to Image Registration

While the superimposition of functional and anatomi- cal images was occasionally attempted, it was during the late eighties that the importance of directly combining anatomy and function began to emerge. The first such attempts were software based [2, 3]. These software de- velopments were driven primarily by a demand for ac- curate localization of cerebral function visualized in PET studies where the low-resolution morphology is, in most cases, insufficient to accurately identify specific cerebral structures. Software fusion techniques were successful for the brain, a rigid organ fixed within the

skull, whereas for other parts of the body, image fusion was found to be somewhat problematic [4, 5].

In particular, combining complementary whole- body image data sets retrospectively is not straight- forward due to the difference in patient set-up and variable deﬁnitions of the axial examination ranges of the two imaging modalities, which are often acquired independently by different medical personal. Further, normal variants of the position and metabolic activity of bowel and intestines at the time of the two scans, as well as dissimilar breathing patterns contribute to ad- ditional systematic difference in the two data sets.

Although some of the positioning errors may be over- come by non-linear image warping techniques and 3D elastic transformations [6] these registration algo- rithms are typically limited to a single anatomical region like the thorax [5-9] and are often labour- intensive, thus making them less attractive for routine clinical use in high-throughput situations.

Hardware Approaches to Combined Imaging

An alternative to post hoc image fusion by software is, instead, to fuse the hardware from the two imaging modalities. While presenting a signiﬁcant number of challenges, such an approach overcomes many of the difﬁculties of the software fusion methods.

In the early nineties, Hasegawa and co-workers at the University of San Francisco developed the first device that could acquire both, anatomical (CT) and functional (SPECT) images, using a single, high-purity germanium detector for both modalities [10, 11]. The CT images were, in addition, used to provide attenua- tion factors for correction of the SPECT data [12], and operating the device with two different energy windows allowed simultaneous emission-transmission acquisitions to be performed. This pioneering work of Hasegawa, Lang and co-workers is important because it was one of the first to take an alternative, hardware- based approach to image fusion. However, the difficulty of achieving an adequate level of performance for both SPECT and CT with the same detector material and without compromising either modality, led the group to explore a combination of SPECT and CT using dif- ferent, dedicated imaging systems for each modality – a clinical SPECT camera in tandem with a clinical CT scanner [13]. The CT images are used to correct the SPECT data for photon attenuation, and the device has been used clinically for patient studies since around 1996.

180 Positron Emission Tomography

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Recognizing the advantages of a hardware solution to combining anatomy and function, Townsend, Nutt and co-workers initiated a design for a combined PET and CT scanner in the early 1990s. The device com- prised a clinical CT scanner and a dedicated clinical PET scanner. As with the CT/SPECT scanner of Hasegawa, the CT images are also used to correct the emission data for photon attenuation. The device was completed in early 1998 and underwent an extensive, three-year clinical evaluation programme at the University of Pittsburgh.

A third possibility to have anatomy and function imaged in a single device is to combine PET and MR.

Obviously such a combination is technologically more challenging than combining PET with CT in view of the extensive restrictions placed on the imaging envi- ronment by the strong magnetic ﬁeld. Nevertheless, proposals to place PET detectors inside an MR scanner also date back to the mid-nineties [14, 15]. In 1996, an MR-compatible PET scanner was developed at UCLA [16], and then in 1997 a second, larger prototype (5.6 cm diameter ring) was constructed [17] and used in collaboration with researchers at Kings’ College London for phantom and animal studies. The studies clearly demonstrated that simultaneous PET and MR images could be acquired using a range of pulse se- quences and at different ﬁeld strengths [18]. The device opened up the possibility of simultaneously imaging [

¹⁸

F]-FDG uptake and measuring MR spectra.

Currently a larger, 11.2 cm PET detector ring is being developed, designed to ﬁt inside a 20 cm diameter magnet bore [19]. However, scaling the design up to human dimensions will present many challenges and is still a number of years away.

Design Concept of the Prototype PET/CT Scanner

Design Concept

The design objective for combined PET/CT imaging, as a technically straightforward combination of two com- plementary imaging modalities, was to provide clinical CT and clinical PET imaging capability within a single, integrated scanner (Table 8.1). The short CT scan dura- tion compared with a typical whole-body PET acquisi- tion time essentially eliminates the requirement for simultaneous CT and PET acquisition. The integration of the two modalities within a single gantry is more straightforward when the simultaneous operation of the CT and PET imaging systems is not required. The PET and CT components are mounted on the same alu- minium support with the CT on the front and the PET at the back, as shown schematically in Fig. 8.2. The entire assembly rotates at 30 rpm and is housed within a single gantry of dimensions 170 cm wide and 168 cm high. The patient port is 60 cm in diameter with an overall tunnel length of 110 cm and a 60 cm axial dis- placement between the center of the CT and the center of the PET imaging ﬁelds. A single patient bed is used for both modalities with an axial travel sufﬁcient to cover 100 cm of combined CT and PET imaging. To simplify the prototype development, the acquisition and reconstruction paths are not integrated, with CT and PET scanning controlled from separate consoles, as shown schematically in Fig. 8.2. Once acquired and reconstructed, the CT images are transferred to the

Table 88.1. A comparison of advantages and disadvantages of software- and hardware-based image fusion.

Software fusion Hardware fusion

Image retrieval from different archives Images avilable from one device

Repeated patient positioning Single patient positioning

Different scanner bed profiles One bed for both scans

Uncontrolled internal organ movement between scans Consecutive scans with little internal organ movement in between Disease progression in time between exams Scans acquired close in time

Limited registration accuracy Improved registration accuracy

Less convenient for patient (two exams) Single, integrated exam

Labour-intensive registration algorithms No further image alignment required

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PET computer to provide the attenuation correction factors for the PET emission data. Final PET recon- struction and CT and PET fused image display is per- formed on the PET computer console.

CT Scanner

The CT components of the prototype PET/CT is a Siemens Somatom AR.SP, a single-slice spiral CT scanner with a 25 kW M-CT 141 tube that produces X- ray spectra of 110 kV

_p

and 130 kV

_p

. The detector array comprises 512 xenon gas-ﬁlled chambers. A slice thick- ness of 1, 2, 3, 5 or 10 mm can be selected, with a maximum rotation speed of 1.3 s. For compatibility with the PET components mounted on the same support, the combined assembly is limited to a rotation speed of 30 rpm. The transverse ﬁeld of view is 45 cm and the patient port is 60 cm. CT data transfer is over mechani- cal slip rings, as is the power to the X-ray tube and de- tectors. The tilting capability of the CT is disabled.

PET Scanner

The PET components mounted on the rear of the support are those of a standard Siemens ECAT

^®

ART scanner [20] comprising dual arrays of bismuth ger- manate (BGO) block detectors. Each array consists of 11 blocks (transverse) by 3 blocks (axial); the blocks are 54 mm × 54 mm × 20 mm in size, cut into 8 × 8 crystals each of dimension 6.75 mm × 6.75 mm × 20 mm. The transverse ﬁeld-of-view (FOV) is 60 cm and the axial FOV is 16.2 cm, subdivided into 24 (partial) rings of detectors for a plane spacing of 3.375 mm.

Shielding from out-of-ﬁeld activity is provided by arcs of lead, 2.5 cm thick, mounted on both sides of the de- tector assembly and projecting 8.5 cm into the FOV beyond the front face of the detectors. The PET scanner

has no septa and the detector arrays rotate continu- ously at 30 rpm to collect the full set of projections re- quired for image reconstruction. Power and serial communications to the rotating assembly are transmit- ted over mechanical slip rings, while high speed digital data transfer is by optical transmission.

Physical performance of the prototype PET/CT

The overall physical performance of the combined scanner is comparable to that of the individual compo- nents, the Somatom AR.SP and the ECAT ART. The PET and CT components are mounted on opposite sides of the aluminium support thus minimizing potential in- terference between the two imaging systems. Although the sensitivity of the PET detectors is temperature-de- pendent, no significant effect from the operation of the X-ray source has been observed. The PET detectors are never exposed directly to the X-ray flux and the opera- tion of the CT has no residual effect on the photomulti- plier tube gains. The PET components can be operated immediately after the acquisition of the CT scan without requiring a recovery time. However, the PET and CT components cannot acquire data simultane- ously because of the high flux of scattered CT photons incident on the PET detectors that results in high levels of random coincidences and system dead-time (from pulse pile-up effects). In view of the short time re- quired for the CT, simultaneous operation of both PET and CT scanners is not considered necessary. The PET components are operated with a detector block inte- gration time of 384 ns [21], a coincidence window of 12 ns, and a lower energy threshold of 350 keV. The operation of the Somatom AR.SP is in accordance with standard CT procedures. Complete details of the results of the performance measurements and relevant para- meter settings can be found in [22].

Figure 88.2. Design of the prototype PET/CT.

The PET components were mounted on the rear of a common rotating support. The axial separation of the two imaging fields was 60 cm.

The entire assembly within the gantry rotated at 30 rpm. The co-scan range for acquiring both PET and CT was 100 cm (maximum). CT and PET scans were acquired and reconstructed on separate consoles but image fusion display was installed on the PET console alone.

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Clinical Protocols and Evaluation

When imaging clinically with the PET/CT, a typical ac- quisition protocol begins with a 260 MBq injection of FDG, followed by a 60 min uptake period. The patient is then positioned in the scanner with the ﬁrst trans- verse section to be imaged aligned with the CT ﬁeld-of- view. An initial scout scan (topogram) is performed to determine the appropriate axial range for the study.

The maximum axial extent of a single spiral scan depends on the deﬁned slice-width and pitch. The total axial length to be scanned is subdivided into contigu- ous, overlapping, 15 cm segments. For the Somatom AR.SP, the spiral scan of each segment typically takes about 40 s, and the 25 kW X-ray tube may sometimes require cooling between segments (Fig. 8.3). Patients are instructed to breathe in a shallow manner during the CT scan. The time for the complete whole-body CT scan is about 5 min. Once the spiral CT covering the re- quired axial length is completed, the patient bed is moved automatically to the start position of the multi- bed PET acquisition, and the PET scan is initiated. An emission scan time of 6-10 min per bed position is se- lected depending on the number of bed positions, re- sulting in a total PET scan duration of 45-50 min (Fig.

8.3). An axial overlap of 4 cm is used between bed posi- tions. The CT images are used for attenuation correc- tion as will be described below, and the corrected emission data are reconstructed using Fourier rebin- ning and attenuation-weighted ordered-subset EM [23]. In this implementation reconstruction takes over one hour to complete.

From July 1998 to July 2001, over 300 patients with a wide variety of different cancers were scanned on the prototype PET/CT [24, 25]. The main indications, most suited to anatomical and functional imaging, are head and neck cancer, and abdominal and pelvic disease, particularly ovarian and cervical cancer. Combined PET/CT in the head and neck is important because normal uptake of FDG in muscles and glands makes interpretation of the studies especially difﬁcult. PET

applications in the abdomen and pelvis are compli- cated by benign, non-speciﬁc uptake in the stomach, intestines and bowel that may be difﬁcult to distin- guish from malignant disease. Combined imaging for clinical routine allows accurate localization of lesions, the distinction of normal FDG uptake from pathology, and the assessment of response to therapy (Fig. 8.4).

Additionally, the use of registered CT and PET images was envisaged for efﬁcient radiation therapy planning, traditionally based on CT alone [26].

CT-based Attenuation Correction

Transforming Attenuation Coefficients

In addition to acquiring co-registered anatomical and functional images, a further advantage of the com- bined PET/CT scanner is the potential to use the CT images for attenuation correction of the PET emission data, eliminating the need for a separate, lengthy PET transmission scan. The use of the CT scan for attenua- tion correction not only reduces whole-body scan times by at least 30% [27], but provides essentially noiseless attenuation correction factors compared to those from a standard PET transmission scan. CT- based attenuation values are, however, energy depen- dent, and hence the correction factors derived from a CT scan at a mean photon energy of 70 keV must be scaled to the PET energy of 511 keV, for which a hybrid scaling algorithm was developed [28]. This scaling ap- proach is based on previous work by La Croix and Tang [29, 30] who have shown that attenuation correc- tion factors for SPECT emission data can be derived from complementary CT transmission data. The single scale-factor approach works well for soft tissues, but serious overestimation of the attenuation properties of cortical bone and ribs is observed, especially with in- creasing difference of the transmission and emission

Figure 88.3. Whole-body FDG acquisition

protocol for the prototype PET/CT. Note, the standard transmission is replaced with a multi- bed CT scan (tube cooling is needed in case of an extended imaging range) and both, IV and oral CT contrast is given to enhance the diagnostic quality of the transmission images. Attenuation correction factors are calculated on-line but PET images are reconstructed post-acquisition.

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photon energies [29]. Nevertheless, the overestimation of attenuation in bone translated into only a minor average overestimation of the tracer uptake in the cor- rected SPECT images due to the low fraction of voxels containing bone compared to other tissues.

In anticipation of the prototype PET/CT the original scaling approach was extended to CT-based attenua- tion correction of PET emission data [28]. A bi-linear scaling is employed to account for both the photon energy difference between CT and PET, and the differ- ent attenuation properties of low-Z (soft tissues) and high-Z (bone) materials in the range of the lower energy X-ray photons (Fig. 8.5).

The algorithm is based on the observation that for water, lung, fat, muscle and other soft tissues, the mass attenuation coefficient (linear attenuation coefficient divided by density) at CT and PET energies is approxi- mately the same. While the actual value is different at the effective CT energy of 70 keV to that at 511 keV, all the tissues can be scaled with a single factor: the ratio of the mass attenuation coefficient at 511 keV to that at 70 keV. The exception is bone because, at CT ener- gies, the mass attenuation coefficient is somewhat higher than that for the other tissues due to the in- creased photoelectric contribution from calcium. The ratio of the mass attenuation at 511 keV to that at 70 keV is approximately 0.53 for soft tissues, and 0.44 for bone.

CT-based PET attenuation correction factors are generated in a 4-step procedure [28]:

1. the CT images are divided into regions of pixels classiﬁed as either non-bone or bone by simple thresholding. A threshold at 300 Hounsﬁeld Units (HU) separates spongiosa and cortical bone from other tissues,

2. the pixel values in the CT image in HU are con- verted to attenuation coefﬁcients of tissue ( μ

T

) at the effective CT energy (~70 keV) using the expres- sion: μ

T

= μ

W

(HU/1000+1); μ

W

is the attenuation coefﬁcient for water,

3. the non-bone classiﬁed pixel values are then scaled with a single factor of 0.53, and bone classiﬁed pixel values are scaled with the smaller scaling factor of 0.44,

4. attenuation correction factors are generated by integrating (forward projecting) along coincidence lines-of-response through the segmented and scaled CT images, with the CT spatial resolution degraded to match that of the PET. Oblique lines-of-response are obtained in the same way by integration through the CT volume.

Today bi-linear scaling methods [28, 31] are widely accepted for clinical PET/CT imaging, and are, with minor modiﬁcations, used routinely for CT-based at- tenuation correction of the PET emission data [32, 33].

Figure 88.4. Male patient with metastatic melanoma before (a) and after (b) chemotherapy in 11/98 and 12/98, respectively. PET/CT images are from the prototype PET/CT. Multiple lesions are depicted as FDG avid and localized accurately within the anatomy of the patient. Each whole-body scan took about 1 h.

PET/CT images before and after therapy are registered by hand and selected axial views are shown to demonstrate multi-variant response to therapy.

a

b

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The time for the acquisition of the attenuation data for a whole-body study can be reduced to one minute, or less, by using a fast CT scan instead of a lengthy PET transmission measurement. Furthermore CT transmis- sion data acquired in post-injection scenarios are not noticeably affected by the emission activity inside the patient due to the high X-ray photon ﬂux [22].

Therefore corrective data processing as in post-injec- tion PET transmission imaging [34, 35] is not required.

While, in principle, the CT-based attenuation coefﬁcients are unbiased and essentially noiseless, there are a number of practical limitations. These include respiration effects, truncation of the CT ﬁeld- of-view when imaging with the arms down, and the effect of using CT contrast.

Patient Respiration

Clinical CT scans of the thorax are normally acquired with breath-hold at full inspiration. The PET image, on the other hand, represents an average over the scan duration of several minutes per bed position, during which the patient breathes normally. Under such a protocol, exact alignment of the CT and PET images, particularly in the lower lungs, is not possible.

Typically the movement of the chest wall is suppressed

by breath holding during the CT scan, and, with the

lungs fully inﬂated, there is a mismatch between the

anterior wall position on CT and the average position

in the PET image, as shown in Fig. 8.6. Incorrect atten-

uation correction factors are then generated by the al-

Figure 88.5. Mass attenuation coefficients for soft tissue and bone (a) differ significantly for lower photon energies for which the photoelectric effect is the dominant interaction with matter. The bi-linear scaling and segmentation approach (b) accounts for the different attenuation properties of soft tissue and bone by, first, segmenting (Se) the CT, and, second, by applying a tissue dependent scale factor (Sc) to these pixels. The greatly increased photon flux used in CT results in essentially noiseless attenuation maps (c) and in attenuation correction factors compared with those derived from a standard PET transmission scans (d).

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gorithm described above, since it is based on the as- sumption that the PET and CT images are accurately co-registered.

When the patient is allowed to breathe normally during both the PET and the CT scans, artefacts appear near the base of the lung and the diaphragm on the CT image (Fig. 8.6b). An alternative protocol is to allow the patient to breathe in a shallow manner (tidal breath- ing) during both the CT and PET scans, a procedure that minimizes the breathing artefacts on CT (Fig. 8.6c) and the mismatch between the PET and CT images.

Nevertheless, a deﬁnitive solution to the respiration problem has yet to emerge.

Truncation of the CT Field-of-View

Clinical CT scans of the thorax and abdomen take a few minutes to acquire and can generally be performed

with the patient’s arms out of the ﬁeld-of-view.

However, in PET/CT imaging total examination time is defined primarily by the time for the emission scan. A typical whole-body PET scan with the prototype PET/CT could last for 1 h, and for scan times of this duration it is difficult for patients to keep their arms comfortably above their head, out of the field-of-view.

More recently, with the introduction of full-ring PET components and faster PET detectors into combined PET/CT designs, total examination time for PET/CT is reduced to 30 min [36], or less. Despite the dramatic reduction in total imaging time some patients may still not tolerate having their their arms raised and sup- ported for the duration of the PET/CT scan, and there- fore CT and PET imaging must be performed with the arms down and close to the body. However, since the transverse ﬁeld-of-view of the CT is 50 cm in diameter (45 cm in the prototype), a small angular range of pro- jections around the anterior-posterior direction is, for

Figure 88.6. The effect of patient respiration on the CT-based attenuation procedure (a). A CT scan acquired with inspiration breath-hold (left) is not matched with the PET image (right) that is acquired with regular breathing. A CT acquired with normal (b) and shallow (c) breathing; the artifacts near the base of the lung (arrow heads) can be reduced (c).

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many patients, truncated (Fig. 8.7). The artefacts caused by truncation affect not only the CT images but also the accuracy of the attenuation correction factors generated from the CT images. The effect is illustrated in Fig. 8.7 for a patient who was imaged on the proto- type PET/CT with arms down. The transverse CT field- of-view is limited to 45 cm in diameter. As shown in Fig. 8.7a, truncation leads to ring artefacts around the arms that affect both the accuracy of the CT images and the attenuation correction factors. As a result the tracer distribution in the reconstructed and corrected emission images appears masked near the arms. Figure 8.7b shows a similar patient study from a second gen- eration PET/CT system. Although the transverse CT field-of-view is increased to 50 cm truncation may still occur when imaging large patients. The theory for an effective correction of these truncation artefacts exists today [37, 38], and simplified correction schemes for

application in the context of PET/CT imaging are cur- rently being pursued (see Chapter 5).

Effects of CT Contrast Agents

Clinical CT scans are acquired with intravenous con- trast and/or oral contrast to enhance the visualization of structures such as the vascular system or the diges- tive tract. CT contrast media use high atomic number substances such as iodine to increase the attenuation of the vessels or the bowel and intestines above normal, non-enhanced values. Depending on the con- centration of the contrast agent, enhancements in CT attenuation values of 1000 HU can be observed. In the presence of contrast agents the routine CT-based atten- uation correction algorithm [28] will incorrectly segment and scale the enhanced structures above

Figure 88.7. Truncation artifacts on CT occur when patients, particularly large patients, are positioned with arms down. The regions outside the maximum transverse field-of-view (45 cm for the prototype PET/CT (a), and 50 cm for second generation PET/CT (b)) are truncated, and the resulting attenuation correction factors for projections traversing the truncated area are underestimated, which yields a “masking effect” of the corrected PET images (CT-AC). The uncor- rected emission images (Em) are shown in the middle row. The transverse field-of-view (fov) of the CT and the PET is indicated by the set of the vertical lines.

Recently, algorithms have become available that help extrapolate the truncated attenuation information (c), which becomes available for CT-based attenuation correction. (Data processing courtesy of Otto Sembritzki, Siemens Medical Solutions, Forchheim, Germany)

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0 HU, resulting in a bias in the attenuation factors. Such biases could potentially generate artefacts in the cor- rected PET images since the contrast-enhanced CT scan is acquired before the PET scan and then used to generate the attenuation coefﬁcients [39, 40].

Intravenous contrast appears in the vessels as focal regions of elevated attenuation on the CT scan, fre- quently with attenuation values above 300 HU.

Although these small regions are scaled as bone, once the CT resolution has been degraded to match that of the PET (step 4 in the algorithm), there is a signiﬁcant reduction in image contrast owing to the effect of smoothing. Pixels associated with the vessels neverthe- less do have a slightly enhanced value due to the pres- ence of contrast even in the resolution-matched CT images, although the overall effect on the attenuation correction factors is found to be negligible.

Positive oral contrast is potentially more problem- atic as it collects in larger-volume structures (e.g., in-

testines) and in a wider range of concentrations. At oral administration, the concentration of the solution will correspond to a pixel value of about 200 HU in the image. However, as water is absorbed from the solution during passage from the stomach and through the in- testines, the concentration increases to corresponding CT values of up to 800 HU. Despite these high CT values, phantom studies show only a 2% increase in the linear attenuation coefﬁcient at 511 keV compared to the value for water. Structures containing oral contrast should therefore be transformed, as they would be in the absence of contrast.

Figure 8.8 shows a transverse CT section through the pelvic region containing both regions of bone and posi- tive oral contrast. In Fig. 8.8a, the histogram of CT pixel values exhibits a plateau, or shoulder, due to both bone and oral contrast enhancement. When setting the pixels containing contrast to 0 HU (Fig. 8.8b) a peak at the origin is generated, and a reduced shoulder that repre-

Figure 88.8. Axial CT image in the presence of positive oral contrast (a) and histogram of CT attenuation values (HU). The same image with the contrast enhanced pixel values set to 0 HU is shown in (b), and for the rectangular region-of-interest containing no bone pixels in (c). Selected histograms without (A) and with (B) the enhanced pixels set to 0 HU. (Reproduced from Valk PE, Bailey DL, Townsend DW, Maisey MN. Positron Emission Tomography: Basic Science and Clinical Practice. Springer-Verlag London Ltd 2003, p.206.).

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sents bone pixels only. Conversely, within the rectangu- lar region-of-interest indicated (Fig. 8.8c) containing contrast-enhanced pixels and no bone, oral contrast is present in pixels with values above about 150 HU as de- termined from anatomy-based segmentation of the en- hanced colon. Thus, to account for the presence of oral contrast in varying concentrations, pixel values above 150 HU should be set to 0 HU before scaling. Above 300 HU, however, a more complex segmentation proce- dure than simple thresholding must be used to distin- guish contrast enhancement from bone. The presence of both intravenous and oral contrast in the same CT section may further complicate this procedure.

Design Concept of a Production PET/CT Scanner

The somewhat-unanticipated demand for combined PET/CT imaging technology that was created to a large extent by the results from the prototype generated a re- sponse from major vendors of medical imaging equip- ment. However, given the choices for the CT and PET components, a number of decisions had to be made (Table 8.2) that included the appropriate level of CT and PET performance, the extent of hard- and software integration, the potential for upgrades, the targeted users and applications, and of course, the cost [41].

Speciﬁcally, the main design questions are:

●

what is an appropriate level of CT and PET perfor- mance?

●

should standard PET transmission sources be provided in addition to CT-based attenuation correction?

●

what is the level of hardware integration that can be achieved?

●

what co-scan range of PET and CT should the patient handling system offer?

●

what level of software integration can be achieved?

In the following sections we describe the design of the ﬁrst commercial PET/CT tomograph that was de- veloped by CTI PET Systems (CPS Innovations, Knoxville, TN) and presented at the Society of Nuclear Medicine Meeting 2000. Since then several other com- mercial PET/CT tomographs from other manufactur- ers have emerged and will be discussed towards the end of this chapter.

The CT and PET Components

The choice of the level of CT and PET performance depends to some extent on the applications envisaged.

As with the prototype, the design described here is tar- geted primarily at PET whole-body oncology, although potential cardiac applications are not excluded.

Since the PET scanner performance is the limiting factor in terms of statistical image quality, spatial reso- lution, and scan duration, the highest possible PET per- formance is obviously indicated. This consideration inﬂuenced the selection of the ECAT EXACT HR+ [42]

as the PET component of choice because of its high sensitivity and high spatial resolution. The scanner has 32 detector rings of crystals of dimensions 4.05 × 4.39

× 30 mm

³

, giving an axial plane spacing of 2.43 mm.

The detectors cover an axial ﬁeld-of-view of 15.5 cm.

The transmission rod sources are removed and all at- tenuation correction factors are derived from the CT images, as described in this chapter. The septa are also removed, resulting in a dedicated 3D PET scanner.

The role of the CT is to provide an anatomical infra- structure for the functional images and accurate atten- uation correction factors for the PET emission data.

However, an objective of this design is also to provide a state-of-the-art CT scan of clinical diagnostic quality, suggesting that a mid to upper range CT scanner can satisfy all three design criteria. For these reasons, the Siemens Somatom Emotion spiral CT scanner (Siemens Medical Solutions, Forchheim, Germany) was selected as the CT component. This CT scanner is available with a single or dual row of Ultra Fast Ceramic (UFC™)

Table 88.2. Design considerations for commercial PET/CT tomographs based on the experiences gained with the prototype PET/CT.

Prototype design Consequence PET/CT design goals

ECAT ART Low-end dedicated PET Highest PET performance

Somatom AR.SP Early 90’s technology High-performance CT

30 rpm Not state-of-art CT Sub-second rotation

60 cm patient port Limitation for RTP Increased port diameter

45 cm CT FOV CT truncation artifacts Increased CT FOV

100 cm co-scan Limitation for whole-body Whole-body co-scan

No bed support Bed deflection possible No bed deflection

Limited access Service difficulties Full service access

Dual acquisitions Operation not integrated Integrated acquisition

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detectors, where each detector row comprises 672 indi- vidual elements. The X-ray tube is a 40 kW Siemens DURA 352 with ﬂying spot technology. The full rota- tion time can be selected from 0.8 s, 1.0 s, and 1.5 s, and slice widths of 1, 2, 3, 5, 8 and 10 mm for the single-slice and 2×1, 2×1.5, 2×2.5, 2×4, 2×5, 8 and 10 mm for the dual-slice. The longest acquisition in a single spiral is 100 s without intermittent tube cooling. The useful di- ameter of the transverse ﬁeld-of-view is 50 cm, and the patient port is 70 cm.

Gantry and Patient Handling System

The gantry dimensions of the Somatom Emotion closely match those of the ECAT EXACT HR+, facilitat- ing the mechanical integration of the two units. In comparison to the original prototype design mechani- cal and thermal isolation is maintained between the two devices for operational and servicing reasons. A schematic of the gantry is shown in Fig. 8.9. The gantry is 188 cm high and 228 cm in width. The overall length is 158 cm, although with the front and rear contouring, the effective tunnel length is only 110 cm. The axial separation of the centres of the CT and PET ﬁelds-of- view is 80 cm. The actual 56 cm patient port of the HR+ is increased to 70 cm to match that of the CT by cutting back the side shielding of the PET. The result- ing patient port diameter is 70 cm throughout the length of the tunnel, which is essential when position- ing most patients from radiation therapy, and which minimizes claustrophobic effects despite the 110 cm tunnel length.

For servicing, the gantries can be separated by moving the PET backwards on rails by about 1 m;

access to the rear of the Emotion CT and the front of the EXACT HR+ is then possible. No service proce- dures on either device have been signiﬁcantly modiﬁed

as a consequence of the particular integration into the PET/CT.

The problem of the increasing vertical bed deflection with increasing distance into the scanner encountered with the prototype was resolved by a complete redesign of the patient handling system (PHS). Support of the patient bed throughout the scan range is important to avoid an increasing vertical deflection of the pallet: the pallet deflects downwards with the patient load, a deflection which increases as the bed moves into the tunnel, adversely affecting the CT and PET image reg- istration accuracy. In the prototype (Fig. 8.10a), the pallet is not supported beyond the cantilever point and an approximate correction for the increasing down- ward deflection is applied in software, based on the patient weight and the pallet position. In the new design, shown in Fig. 8.10b, a carbon fibre pallet is sup- ported at one end by a pedestal that moves horizontally on floor-mounted rails driven by a linear motor. Since the cantilever point does not change, the vertical deflection is limited to a few millimetres once the patient is aligned on the bed, allowing sub-millimeter intrinsic registration accuracy to be achieved between the CT and the PET, independently of the patient weight. A total length, including the head holder, of 145 cm can be scanned with both CT and PET. A flat pallet option is available for use with the PHS when scanning patients undergoing PET/CT for radiation therapy treatment planning.

Software Integration

A key feature of the production PET/CT scanner com- pared to the prototype is the integration of the CT and PET acquisition and reconstruction software on a single console within a modality-independent software

Figure 88.9. A schematic of the biograph PET/CT scanner, as a representative example of the second generation, commercial PET/CT systems. The axial separation of the CT and PET imaging field is 80 cm. A common cantilever design patient handling system (PHS) is mounted to the front of the combined gantry for accurate patient positioning across the 145 cm co-scan range.

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environment (syngo

^®

, Siemens Medical Solutions, Erlangen, Germany). The CT and PET scans can be ac- quired, reconstructed and viewed on a single console by selecting the appropriate syngo task card. An example of the PET/CT examination cards is shown in Fig. 8.11a. The reconstruction software includes CT- based attenuation correction, Fourier rebinning and an attenuation-weighted ordered-subset EM algorithm [43]. The complete whole-body, attenuation-corrected PET images are available within a couple of minutes of the completion of the scan, and all image formats are DICOM compliant to facilitate transfer to PACS or radi- ation therapy planning systems.

The routine availability of registered CT and PET images highlights the importance of a fused image viewer with a full set of features (Fig. 8.11b). These include transverse, coronal and sagittal displays of CT, PET and fused images using an alpha-blending fusion algorithm. Each modality has the usual set of speciﬁc features, for example preset windows and measure- ment tools for CT, and region-of-interest (ROI) manip- ulation and SUV calculations for PET. To take full advantage of the registered data sets, enhanced viewing features are required such as linked cursors and common ROI and measurement tools on both CT and PET.

Physical Performance

Other than eliminating the tilt option, no major modifications were required to integrate the Siemens Emotion CT scanner into the PET/CT. The perfor- mance characteristics are therefore identical to a stan- dard Emotion CT scanner. The EXACT HR+, however, was modified to accommodate a 70 cm patient port by cutting back the side shielding, and in the absence of septa, all operation is in 3D mode. The increased patient port does not have a significant effect on

scanner performance unless there are high levels of ac- tivity outside the ﬁeld-of-view, as can be the case with whole-body imaging. In this situation, with 3D only, the reduced side-shielding results in an increased level of single and scattered photons incident on the detec- tors, increasing the randoms rate and lowering the peak noise equivalent count rate (NEC). When oper- ated with a 384 ns integration time, the peak NEC for the biograph measured with a 70 cm phantom accord- ing to the NEMA 2001 standards is 20-30% less than that of a standard EXACT HR+ operated with a 768 ns block integration time. However, if the random coinci- dences are smoothed prior to subtraction from the prompts (see Chapter 6), the peak NEC for the bio-

graph is comparable to that of a standard HR+. Hence,

by using a shorter integration time, and implementing random coincidence smoothing, the reduced side- shielding in the biograph results in no appreciable degradation in the performance of the PET component compared to a standard EXACT HR+.

State-of-the-Art PET/CT Systems

Second Generation

The ﬁrst commercial scanners (Table 8.3) that followed successful clinical imaging with the prototype PET/CT [24, 25, 44, 45] appeared in early 2001 and consisted of a CT scanner in tandem with a PET scanner, with little or no mechanical integration of the two systems beyond a common gantry cover [46]. This approach, with the CT and PET scanners kept separate, has been a characteristic feature of all commercial designs (Fig.

8.12). The advantage of this design is ﬂexibility in that different levels of CT and PET performance can be combined as required and upgraded independently. An upgrade path from CT or PET to PET/CT can, in some cases, also be envisaged.

Figure 88.10. A vertical offset (arrows) between the CT and PET data of the same exam is introduced when using a patient handling system with a centre of gravity fixed with respect to the gantry (a). The relative vertical offset between the CT and PET data is eliminated when translating the entire patient handling system, i.e., the patient is moved together with the support structure into the gantry (b).

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Figure 88.11. (a)State-of-the-art PET/CT tomographs offer joint acquisition and data processing consoles. A single console is used to register patient information, select and acquire a pre-defined acquisition protocol, and reconstruct image sets. (b)Selective screenshots of commercial PET/CT fusion display options: (i) Reveal MVS from CTI Molecular Imaging Inc/Mirada Solutions, (ii) Xeleris by GE Medical Systems, (iii) Syntegra by Philips Medical, and (iv) e.soft 3.0 by Siemens Medical Solutions.

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All designs (Table 8.3) incorporate a common patient couch that is designed to eliminate or minimize vertical deflection due to the weight of the patient and to ensure accurate alignment of the CT and PET images (Fig. 8.10). The actual approach to address this issue, however, varies among the manufacturers of PET/CT systems. In one design vertical deflection of the pallet is eliminated through a pedestal-based patient handling system (Fig. 8.10b), while in another design a standard patient bed is mounted on a rail system in the floor to bridge the distance between the CT and the PET field-of-view. Alternatively, the bed support system is fixed with respect to the gantry, and the pallet moves on rails inside the combined system where it is supported further by a post. The maximum co-axial imaging range with any of these patient han- dling systems is 145 cm to 200 cm.

With the detector and data acquisition sub-systems kept separate in the second generation commercial PET/CT systems, attempts have been made to integrate the acquisition and data processing more closely.

Combined scanners are operated from a single console with application-speciﬁc task cards selected for the CT and PET acquisition (see Fig. 8.11). While the early designs involved multiple computers for acquisition, image reconstruction, and image display (Fig. 8.2), progress is being made in combining some of these functions into one computer, thus reducing complexity and increasing reliability. The most recent PET/CT systems are simpler to operate, involve fewer computer systems, and are considerably more reliable. Indeed, poor reliability would be a major concern for the high patient throughput attainable with these scanners.

As mentioned, PET/CT designs have taken advantage of recent advances in both CT and PET technology by maintaining separation of the imaging sub-systems.

While the CT scanners in the ﬁrst PET/CT designs

were single or dual-row, more recently 4-, 8- and even

16-row systems have been incorporated into PET/CT

scanners. These CT modules offer sequential as well as

spiral scanning modes with increased X-ray tube heat

capacities to cover large volumes in short scan times

Figure 88.12. Prototype (a), second generation (b), and third generation (c) PET/CT tomographs with major distinguishing features regarding CT and PET technology. Second generation PET/CT employ alternative (LSO- and GSO-based) PET technology for faster whole-body imaging, while ultra-fast and multi-row CT technology in third generation PET/CT opens the applications for cardiac imaging and gated therapy planning. Systems shown from left to right: prototype, Discovery LS (GE Medical Systems), biograph-BGO/LSO (Siemens Medical Solutions), Gemini (Philips Medical Systems), Discovery ST (GEMS), biograph Sensation 16 (SMS).

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[47, 48], a necessary prerequisite for oncology imaging when examining patients with only limited breath-hold capabilities. Alternatively, imaging ranges now can be scanned with ﬁner axial sampling in similar or less time than with a single-row CT. Furthermore, when employing IV contrast agents several organs can be imaged at peak contrast enhancement during a single spiral CT, or repeat CT exams can be acquired for a single IV contrast injection over an individual organ, such as the liver [49]. The use of spiral CT technology in combined PET/CT imaging therefore offers high- quality CT images for a variety of imaging conditions encountered in clinical oncology.

The recent developments in PET scanner technology have been primarily oriented towards the introduction of new scintillators. For over two decades, PET detec- tors have been based on either thallium-activated sodium iodide (NaI(Tl)) or bismuth germanate (BGO).

While NaI(Tl) has high light output, it has low stopping power at PET photon energies (511 keV) and a long decay time. The low stopping power was overcome by BGO, ﬁrst introduced for PET in 1977 [50] but at the expense of considerably reduced light output com- pared with NaI(Tl); the decay time for BGO is 30%

longer than NaI(Tl). Some physical properties of these scintillators are compared in Table 8.4. The introduc-

Table 88.3. Design and performance parameters of PET/CT systems. PET Performance parameters were acquired to the NEMA 2001 standard.

(CPS Innovations is a joint venture of Siemens Medical Solutions, Inc and CTI Molecular Imaging, Inc).

GE Medical Systems Philips Medical

PET/CT PET/CT

2000 2002 2002

Generation 22 Generation 33 Generation 22

208 / 203 / 205 192 / 230 / 140 206 / 230 / 590 Height/Width/Depth [cm]

142 cm 100 206 with gap (30) Inner tunnel length [cm]

Tapered, 70 cm–60 cm Uniform, 70 cm 70 cm CT, 63 cm PET Patient port diameter

160 cm 160 cm 195 cm Standard co-scan range

floor-mounted pedestal floor-mounted pedestal floor-mounted, dual pallet PHS

200 kg 200 kg 202 Max patient weight

no yes yes Radiation therapy table attachment

LightSpeed Plus LightSpeed Range Mx8000D System components

yes yes yes Spiral CT

4, 8, 16 4, 8, 16 2, 16 Max number of active detector rings

Solid state–Lumex Solid state - Lumex Solid State Detector material

pixel array pixel array 2D solid state detector array Detector design

50 cm 51 cm 50 cm Measured transverse FOV

0.625 mm–10 mm 0.625 mm–10 mm 0.5, 1, 2.5, 5, 8, 10, 16 mm Min and Max Slice width [mm]

75 rpm 75 rpm 120 rpm Max rotation speed

8.5 lp/cm 8.5 lp/cm 22 lp/cm Maximum spatial Resolution

5 mm / 3 HU / n.a. / 34 mGy / n.a. 6 mm / 3 HU / n.a. / 34 mGy / n.a. 4 mm/ 3 HU/ [n.a.]/ 27mGy/ [n.a.] Detectability

6.8 mGy (120 kVp) 6.8 mGy (120 kVp) 14–28 mGy (120 kVp) Centre dose (CTDI_100, body

phantom) per 100 mAs

ADVANCE Nxi unique, not available separately Allegro System components

2D and 3D 2D and 3D 3D PET Acquisition mode

full-ring full-ring full-ring Detector design

yes yes no Septa

BGO BGO GSO Detector material

optional no ¹³⁷Cs point (740 MBq) Transmission sources

55 cm 70 cm 57.6 mm Transverse FOV

15.2 cm 15.2 cm 18 cm Axial FOV

35 47 90 Transverse images per bed

4.2 mm 3.2 mm 2 mm Image plane separation

6.3 mm (2D), 6.4 mm (3D) 5.2 mm (2D), 5.8 mm (3D) 5.4 mm Axial resolution

4.8 mm (2D), 4.8 mm (3D) 6.2 mm (2D), 6.2 mm (3D) 4.8 mm Transverse Resolution

1.3 cps/kBq (2D), 6.5 cps/kBq (3D) 1.9 cps/kBq (2D), 9 cps/kBq (3D) 3.8 cps/kBq Sensititivity 165 kcps at 130 kBq/mL (2D), 82 kcps at 46 kBq/mL (2D) 45 kcps at 9 kBq/mL Peak NEC 42 kcps at 8.5 kBq/mL (3D) 62 kcps at 9.5 kBq/mL (3D)

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tion of faster scintillators such as gadolinium oxy- orthosilicate (GSO) [51] and lutetium oxyorthosilicate (LSO) [52], also compared in Table 8.4, offer enhanced PET scanner performance. LSO in particular out- performs BGO in almost every aspect, especially light output and decay time. The faster scintillators (Table 8.4) have lower dead time and give better count rate performance, particularly at high activity concentra- tions. This behaviour is conﬁrmed in Fig. 8.13 where the Noise Equivalent Count Rate (NEC) [53] is shown as a function of activity concentration in the NEMA NU- 2001 phantom for 2D and 3D BGO scanners compared with a 3D LSO scanner. The curves show the expected behaviour with 3D being superior to 2D and LSO out- performing BGO even at low activity concentrations.

The new pico-3D electronics, matched to the physical properties of LSO, show a signiﬁcant improvement over the older design with peak NEC exceeding 80 kcps. All 3D measurements are for PET/CT scanners, whereas the 2D data are for the standard ECAT EXACT PET scanner (CPS Innovations, Knoxville, TN) with septa extended.

These advances in PET detector technology are reﬂected in the variety of PET design options in cur- rently available PET/CT technology, which, depending

on the manufacturer, employ BGO-, LSO-, or GSO- based detector technology. In the Discovery LS, for example, the PET tomograph is based on bismuth ger- manate (BGO), a scintillator that is the most widely employed detector material in PET. The ECAT

HR+/Somatom Emotion (Duo) also uses BGO as the

PET detector material of choice. CPS Innovations also offers a PET/CT model (ACCEL/Somatom Emotion

(Duo)) with the PET detector being based on lutetium

oxyorthosilicate [54]. Accepting the potential for using faster crystals in PET imaging technology, Philips Medical Systems offer a combination of the GSO-based

Allegro PET tomograph (GSO: gadolinium oxyorthosil-

icate) with a state-of-the-art CT scanner. Most com- mercial PET/CT designs favour the use of 3D-only emission acquisitions by eliminating the septa from the PET components. Assuming proper data processing, 3D PET offers a number of advantages over 2D PET, such as higher sensitivity and higher count rates at lower ac- tivity concentrations [55]. Further, the sensitivity ad- vantage of 3D imaging and the fast scintillation properties of either GSO or LSO can be combined to result in high-quality PET images at reduced scan times and increased patient comfort.

Third Generation

By introducing 16-ring CT technology into combined PET/CT designs the advantages of very fast and high- resolution volume coverage by CT are translated di- rectly into the context of anato-metabolic imaging.

Durable CT operation at scan speeds of 0.5 s and less help reducing respiration-induced artefacts in ex- tended imaging ranges as well as imaging the anatomy of multiple organs at peak enhancement after intra-

Figure 88.13. Noise Equivalent Count Rate (NEC, k=1) as a function of activity concentration in the NEMA NU-2001, 70 cm phantom. The curves are shown for the standard BGO ECAT EXACT HR+ both, with septa extended (BGO 2D) and with septa retracted (BGO) 3D), a BGO PET/CT scanner (BGO 3D), an LSO PET/CT scanner with standard electronics (LSO 3D), and an LSO PET/CT scanner with the new, high count-rate pico3D electronics (LSO pico3D).

Table 88.4. Physical properties of different scintillators for PET

Property NaI(Tl) BGO LSO GSO

Density [[g/mL] 3.67 7.13 7.4 6.7

Effective ZZ 51 74 66 61

Attenuation llength [[cm-1] 2.88 1.05 1.16 1.43

Decay ttime [[ns] 230 300 35–45 30–60

Photons/MeV 38,000 8,200 28,000 10,000

Light yyield [[% NNaI] 100 15 75 25

Hygroscopic Yes No No No

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venous contrast injection. As multi-slice CT pro- gresses to a greater number of slices and shorter rota- tion times, a critical review of the actual requirements of PET/CT for speciﬁc applications will be necessary.

The beneﬁt of freezing motion during the acquisition of the CT and imaging the anatomy of the patient at a particular point during involuntary periodic motion cycles (cardiac and respiration), for example, is lost without adequate acquisition modes for the PET exam portion. Many of the recent CT improvements are tar- geted at cardiology, whereas the role of PET/CT in cardiology has yet to be established. Indeed, oncology applications may be adequately addressed with a lower performance CT scanner, such as a 2 or 4-row system [56].

Current developments in improved acquisition and data processing software accompany new hardware developments for updated combined tomograph designs: the

Discovery ST from GE Medical Systems,

and a combined LSO-PET/16-ring CT from CPS Innovations (Table 8.3). The ST, for example, is based on a revised full-ring BGO-PET system with 10,080 BGO crystals of 6.3 × 6.3 × 30 mm

³

being arranged in 280 detector blocks (6 by 6) with a somewhat reduced detector ring diameter (88.6 cm) compared to the pre- decessor (Advance Nxi, 92.7 cm ring diameter) for im- proved sensitivity in whole-body imaging [57]. The ST is available with a choice of 4-, 8-, or 16-ring CT tech- nology. The LSO-PET/16-ring CT by CPS Innovations offers similar-size detector crystals at a reduced coinci- dence window time (4.5 ns vs 11.5 ns) and a light

output that is ﬁve time that of BGO (Table 8.4). This PET/CT is available exclusively with 16-row CT.

Optimized Protocols for Routine Clinical Procedures

General Considerations for FDG-PET/CT Imaging Protocols

For oncology purposes a standard PET/CT acquisition protocol, in essence, is a modern-day PET oncology imaging protocol, which consists of three steps: (1) patient preparation and positioning, (2) transmission scan, and (3) emission scan. Additional CT scans, such as, e.g., a 3-phase liver CT, or a high-resolution lung scan could be requested by the reviewing physician, but generally these CT scans are not used for attenua- tion correction. While the clinical acquisition protocols of the PET/CT systems today are similar to those from the prototype [58], demands for high diagnostic image quality increase rapidly, and therefore a number of general and speciﬁc considerations apply to current PET/CT imaging protocols (Fig. 8.14). These considera- tions are described in detail for each of the steps of a combined PET/CT acquisition in [59] and [60]. All PET/CT tomographs offer the use of the available CT transmission images for CT-based attenuation correc-

Figure 88.14. Standard FDG-PET/CT imaging protocol (from left to right). The patient is positioned on a common patient handling system in front of the combined gantry (a). First, a topogram is used to define the co-axial imaging range (a). The spiral CT scan (b) precedes the emission scan (d). The CT images are reconstructed on-line and used for the purpose of automatic attenuation correction of the acquired emission data (c). CT, PET with and without attenuation correction, and fused PET/CT images can be used for the clinical image review (e).

a) b) c) d) e)

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tion. To account for the systematic difference in routine acquisition protocols for CT and PET, and to avoid artefacts on the CT transmission data, which may propagate into the attenuation-corrected emission data, particular attention must be given to the prepara- tion of the patient for the CT and to the CT acquisition.

Modifying CT Acquisition Parameters in PET/CT Imaging

Respiratory Motion

Several PET/CT groups have described respiratory motion and the resulting discrepancy of the spatial in- formation from CT and PET as a source of potential artefacts in corrected emission images after CT-based attenuation correction [58, 61, 62]. These artefacts become dominant when standard full-inspiration breath hold techniques are transferred directly from clinical CT to combined PET/CT examination protocols scanning without suitable adaptations (Fig. 8.6a). In the absence of routinely available respiratory gating options the anatomy of the patient captured during the CT scan must be matched to the PET images that are acquired over the course of multiple breathing cycles. Reasonable registration accuracy can be obtained, for example, with the spiral CT scan being acquired during shallow breathing [61, 63, 64]. Alternatively, a limited breath hold protocol can be adopted with either a 1- or a 2-row system, or when dealing with uncooperative patients.

Patients are then required to hold their breath in expira- tion only for the time that the CT takes to cover the lower lung and liver, which is typically around 15 s [65].

Breath hold commands (in normal expiration, for example) can be combined with very fast CT scanning, and therefore may help reduce respiration mismatches over the entire whole-body examination range. With multi-row CT, such as in third generation PET/CT systems (Table 8.3), it is now possible to scan the entire chest at high resolution within a single breath hold.

Nevertheless, when respiration commands are not tol- erated well and signiﬁcant respiration-induced arte- facts are suspected [66], it is advisable to reconstruct the emission data without attenuation correction and to review the two sets of fused PET/CT images very carefully.

Use of CT Contrast Agents

Clinical CT examinations are almost routinely per- formed with contrast enhancement to selectively in-

crease the visibility of tissues and organs for an easier and more accurate assessment of disease and existing alterations of the anatomy of the patient. To achieve a diagnostic beneﬁt in contrast-enhanced CT a number of – frequently competing – parameters during the ap- plication of the contrast agent must be considered [67].

The unmodiﬁed transfer of standard CT contrast pro-

tocols into the context of PET/CT has been shown to

yield CT and PET images with contrast-related arte-

facts if attenuation correction is performed based on

the acquired CT data [60]. Depending on the contrast

concentration CT-based attenuation coefﬁcients were

overestimated by 26 % [40] up to 66 % [68]. However,

the resulting overestimation of the standardized

uptake values (SUV) in the corresponding regions on

the corrected PET was only 5 % and thus clinically

insigniﬁcant assuming the contrast materials were dis-

tributed evenly [69]. Nevertheless, the concentration of

the oral contrast agent in the colon can vary

signiﬁcantly as reported by Carney et al. [70] and may

lead to a degradation of the diagnostic accuracy of the

corrected PET data. Therefore, threshold-based seg-

mentation algorithms to segment and replace contigu-

ous areas of high-density contrast enhancement on CT

images prior to the attenuation correction procedure

have been developed [40, 71]. For example, Carney

et al. have shown that a modiﬁcation can be made to

the original bi-linear scaling algorithm by Kinahan et

al. [28] to separate contrast-enhanced CT pixels from

those of bone, as shown in Fig. 8.15. Pixel enhancement

from positive oral contrast is around 200 HU at inges-

tion through the stomach, increasing to about 800 HU

in the lower GI tract as water is absorbed from the

contrast solution (Fig. 8.15a). Starting with a contrast

enhanced CT scan, cortical bone pixels are identiﬁed

with a threshold greater than 1500 HU and a region

growing algorithm used to identify all contiguous

pixels with bone content. The skeleton can then be ex-

tracted from the CT images. Contrast-enhanced pixels

are identiﬁed by applying a simple threshold at, for

example, 150 HU, well above any soft tissue value. The

CT image pixels identiﬁed as oral contrast can be set

to a tissue-equivalent value thus ensuring accurate at-

tenuation correction factors for the PET data. Since the

presence of contrast material has a negligible effect at

511 keV, the spatial redistribution of the contrast mate-

rial between the CT and PET scans during the total

imaging time does not create a problem; the aim of the

modiﬁed algorithm is to remove the effect of contrast

from the CT images and avoid incorrect scaling of con-

trast-enhanced regions. The modiﬁed algorithm can, to

a considerable extent, also reduce artefacts due to

catheters and metallic objects in the patient.

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Unlike segmentation techniques that aim at retro- spective modiﬁcations and corrections of the mea- sured CT-based attenuation map alternative contrast application schemes represent a straightforward ap- proach to avoiding artefacts from high concentrations of positive oral (and IV) contrast agents prospectively.

Antoch et al. have presented a water-based oral con- trast agent, resulting from previous developments for improved MRT contrast enhancement, for PET/CT imaging [72]. This contrast agent is based on a combi- nation of water, 2.5% mannitol, and 0.2% of locust bean gum and allows for good differentiation of bowel loops from surrounding structures (Fig. 8.15b). Unlike

iodine, or barium, water-equivalent oral contrast agents do not increase the CT attenuation and thus do not lead to an overestimation of the PET activity in the corrected images.

While alternative contrast materials are inadequate for vascular enhancement, high-density artefacts from the bolus injection of IV contrast agents [39] can be avoided by alternative acquisition protocols [73]. For example, diagnostic quality CT can be achieved and focal contrast enhancement in the thoracic vein can be avoided under the condition of the caudo-cranial (i.e., reverse CT scanning following a somewhat prolonged scan delay after the administration of the IV contrast)

Figure 88.15. Positive oral contrast agents may lead to artifacts on PET/CT images, and thus must be accounted for. For example, CT-based attenuation correction can be modified to account for positive oral contrast agents retrospectively (a). The modified algorithm applies a region-growing technique to the contrast enhanced CT images (left) to extract the skeleton (middle). The contrast-enhanced pixels (right) can then be identified by a simple threshold at 150 HU since after removal of the skeleton the only pixels with values above 150 HU will be those with contrast enhancement. Alternatively, a water-based oral contrast may be used instead of the positive oral contrast (b). Acceptable distention of the bowel and artifact-free fully-diagnostic CT images of the abdomen can be achieved with water-based oral contrast, as seen from the coronal CT (left), transverse CT (middle) and corrected PET (right).

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(Fig. 8.16). A more general solution would be to acquire only a non-enhanced CT for attenuation correction and anatomical labelling. However, additional CT scans with contrast enhancement might be required then for accurate delineation of lesions, thus leading to addi- tional patient exposure and more logistical efforts.

Metal Artefacts

High-density implants, such as dental ﬁllings, pace- makers, prostheses, or chemotherapy infusion ports may lead to serious artefacts in CT images [74, 75].

These CT artefacts have been shown to propagate through CT-based attenuation correction into the cor- rected PET emission images where artiﬁcially in- creased tracer uptake patterns may then be generated [76-78]. It is therefore recommended that PET images from PET/CT are routinely correlated with the comple- mentary CT, and that these PET data are interpreted with care when lesions are observed in close proximity to artefactual structures on CT. Until robust metal arte- fact correction algorithms [75, 79] become available routinely in PET/CT the additional evaluation of the emission data without CT-based attenuation correction is also recommended [78].

Truncation Artefacts

Spiral CT technology currently offers a transverse field-of-view of 50 cm, and thus falls short 10 cm less than the corresponding transverse PET field-of-view (Table 8.3). This difference may lead to truncation arte- facts in the CT images [80] and to a systematic bias of the recovered tracer distribution when scanning obese patients, or when positioning patients with their arms down (Fig. 8.7). If not corrected for truncation, CT images appear to mask the reconstructed emission data with the tracer distribution being only partially recovered outside the measured CT field of view.

To reduce the amount of truncation on CT and to minimize the frequency of these artefacts, whole-body

or thorax patients should be positioned according to CT practice with their arms raised above their head. By keeping the arms outside the ﬁeld-of-view the amount of scatter [80] and patient exposure are also much reduced. Given the short acquisition times of a PET/CT most patients tolerate to be scanned with their arms raised for the duration of the combined exam.

A number of algorithms have been suggested to extend the truncated CT projections and to recover the unmeasured regions of the attenuation map in cases where truncation is observed. If applied to the CT images prior to CT-based attenuation correction these correction algorithms will help to recover completely the tracer distributions measured with the comple- mentary emission data [81]. Further work is needed, however, to make such algorithms routinely available for clinical diagnostics.

Future Perspectives for PET/CT

The trend of PET/CT scanners is perhaps best illus- trated by a design in which a 16-slice CT scanner, the Sensation 16 (Siemens Medical Solutions, Forchheim, Germany) is combined with the recently-announced high-resolution, LSO PET scanner (CPS Innovations, Knoxville, TN). The new PET scanner has unique 13 x 13 LSO block detectors each 4 mm x 4 mm in cross- section (Fig. 8.17). The pico-3D read-out electronics, adapted to the speed and light output of LSO, is oper- ated with a coincidence time window of 4.5 ns and a lower energy threshold of 425 keV. The signiﬁcance of these high-resolution detectors is illustrated in Fig. 8.17 for a patient with squamous cell carcinoma of the right tonsil. Following treatment that included a right tonsillectomy, radical neck dissection and chemotherapy, the patient was restaged by scanning ﬁrst on an ECAT EXACT (CPS Innovations, Knoxville, TN) with 6.4 mm x 6.4 mm BGO detectors, and then on

Figure 88.16. Alternative schemes of

application of IV contrasts agents are being pursued to avoid contrast-induced artifacts on PET/CT images while providing acceptable image quality to the radiologists.

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the high resolution PET/CT scanner. A coronal section from the PET scan demonstrates a diffuse band of ac- tivity in the right neck. The corresponding PET section from the PET/CT scan (Fig. 8.17) resolves this diffuse band into individual nodes in the neck of the patient.

The recent introduction of the fast scintillators LSO and GSO as PET detectors has occurred at just the right moment for PET/CT where a reduction in the lengthy PET imaging time is essential to more closely match that of the CT. These tomographs are aimed primarily at high throughput with whole-body imaging times below 30 min. While it is unlikely that whole- body PET imaging times will be reduced to the 30-60 s that is required for CT scanning, a scan time less than 10 min is feasible with new high-performance LSO area detectors currently under development. Such a design will represent a breakthrough in cancer imaging, eliminating problems of patient movement and trun- cated CT ﬁeld-of-view, and substantially reducing artefacts due to respiration. Throughput will increase signiﬁcantly, as will patient comfort and convenience.

New applications, such as dynamic whole-body scans

and the use of short-lived radioisotopes (e.g.,

¹¹

C with a 20 min half-life) will then be within reach.

Future developments in combined PET/CT scanners will be exciting, attaining a higher level of integration of anatomical and functional imaging performance than before. By fulﬁlling an important role, not only in the diagnosis and staging of cancer, but in designing and monitoring appropriate therapies, the combined PET/CT scanner will undoubtedly have a signiﬁcant impact on patient care strategies, patient survival and quality of life.

Acknowledgements

The combined PET/CT project involved many people at the University of Pittsburgh, CTI PET Systems in Knoxville, Tennessee and Siemens CT division in Forchheim, Germany over the last few years. In partic- ular, we acknowledge the seminal contribution of Dr

Figure 88.17. High-resolution PET and PET/CT imaging using LSO-based detectors (Hi-Rez).

A 52 y/o male patient, 70 kg, diagnosed with squamous cell tonsillar cancer and a 4 cm positive node in the neck. The patient underwent pre-surgical chemotherapy, a right tonsillectomy and a right radical neck dissection for removal of the positive node and 45 additional nodes; all of the additional nodes had negative pathology. The patient suffered post-surgical infectious

complications. A follow-up PET scan (Standard) acquired with arms down showed a diffuse band of activity in the right neck (arrow) seen on a coronal section. A PET/CT scan acquired with arms up and with the new high-resolution LSO-based detector blocks (Hi-Rez) clearly resolved this diffuse band of activity into individual, sub clinical lymph nodes (arrow).