Performance characterization and development of quantitative procedures for PET-CT scanners

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Thesis

Reference

Performance characterization and development of quantitative procedures for PET-CT scanners

SCHOENAHL, Frédéric François

Abstract

Positron Emission Tomography (PET) imaging is a functional exploration technique for humans and animals invented in the early 70s, based on the administration of a specific radioactive labeled molecule. The first large part of this work reports theoretical aspects of the technique, including the problem of image reconstruction and corrections based on physical models necessary for the realization of 3D PET imaging. PET has gained large adoption in the clinics today, especially in the field of oncology. Considering this new availability, the question of standardization arises. As a consequence we explore and compare absolute performances of a unique panel of >30 PET-CT cameras based on a recognized standard.

We write some proposals for further standardization, and for extension to evaluation of a novel commercial feature: time-of-flight imaging. We further develop image processing techniques for spatial resolution recovery based on a general analytic model, adapted to small lesions (MBPVEC). Another related development concerns larger lung lesions when respiratory motion is present. Quantitative corrections based on CT are [...]

SCHOENAHL, Frédéric François. Performance characterization and development of quantitative procedures for PET-CT scanners . Thèse de doctorat : Univ. Genève, 2011, no.

Sc. 4320

URN : urn:nbn:ch:unige-167694

DOI : 10.13097/archive-ouverte/unige:16769

Available at:

http://archive-ouverte.unige.ch/unige:16769

Disclaimer: layout of this document may differ from the published version.

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UNIVERSITÉ DE GENÈVE

Département de Radiologie et FACULTÉ DE MÉDECINE

informatique médicale PD H. Zaidi

Département d’informatique FACULTÉ DES SCIENCES

Pr S. Voloshynovskiy

Performance characterization and development of quantitative procedures for PET-CT scanners

THÈSE

présentée à la Faculté des sciences de l’Université de Genève pour obtenir le grade de Docteur ès sciences, mention interdisciplinaire

par

Frédéric François Schoenahl de

Mulhouse, France

Thèse N°4320 GENÈVE

Service de reprographie de l’Uni Mail 2011

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Résumé

Caractérisation de la performance et développement de procédures quantitatives pour les scanners TEP/TDM

[Performance characterization and development of quantitative procedures for PET-CT scanners]

Thèse de doctorat, Université de Genève, Suisse par Frédéric Schoenahl

La Tomographie par Emission de Positrons (TEP) est une technique d’exploration fonctionnelle inventée au début des années 1970 qui a connu un développement particulièrement important ces 20 dernières années. Le mode d’acquisition d’un appareil de TEP est basé sur la physique particulière du positron, qui détecte les résultats de sa décroissance. La paire de photons d’annihilation résultant de la formation d’un complexe du positron avec un électron de la matière a une corrélation spatiale bien définie et caractéristique. Cette propriété peut être utilisée avec une statistique suffisante pour reconstruire une région de l’espace où la probabilité d’occurrence de l’annihilation est maximale, et donc où la quantité d’émetteur de positrons est susceptible de s’être désintégrée. L’application médicale consiste donc en la construction et l’étude de la distribution spatiale de molécules marquées d’un émetteur de positrons, et de sa représentation spatiale par la modalité et des moyens informatiques. Elle trouve en particulier une application reconnue en oncologie clinique, pour la recherche de tumeurs en utilisant un analogue du glucose le ¹⁸F-fluoro- deoxy-glucose (FDG) et grâce à un nombre croissant de radio-traceurs alternatifs permettant d’autres applications telles que le diagnostic de la perfusion myocardique ou l’étude des fonctions cérébrales. L’étude et le développement de la modalité pour une application clinique demande donc un champ de compétences variées, et sa performance doit être conditionné par le critère d’utilité en médecine, mais aussi par les possibilités importantes en recherche fonctionnelle.

Dans ce travail nous rappelons les développements modernes de la technique d’imagerie, en suivant la dualité entre les processus de traitement de signal de la chaine de détection, et les aspects de modélisation pour le processus général de la reconstruction de la bio-distribution de l’émetteur de positrons. Nous détaillons en particulier les développements nécessaires à l’imagerie TEP en 3D, qui fut un des déclencheurs majeurs pour une efficacité clinique de la technique.

L’acquisition tridimensionnelle avec une collimation limitée pour la détection pose des problèmes généraux et fondamentaux. La modalité prit cependant cette orientation à la fin du XX^ème siècle dû essentiellement à l’important bénéfice en sensibilité et grâce aux possibilités nouvelles de traitement du signal et modélisation de la physique. La forte corrélation spatiale des photons d’annihilation permet un traitement idéalement séquentiel des évènements de détection et la quantité de la molécule investiguée peut être retrouvée. Cet avantage important de la technique requiert cependant une série de corrections dont nous discutons le modèle dans ce travail. Le taux de comptage en TEP est très faible en comparaison d’autres techniques comme le CT, et la gestion du bruit dû à l’atténuation ou la diffusion des photons est un problème statistique complexe qui nécessite la cohabitation de modèles précis au sein d’algorithmes d’estimation. Nous rappelons donc dans la première partie de cette thèse les fondements de la reconstruction 3D basés sur la solution du problème inverse de l’acquisition, et discutons en particulier les méthodes actuelles d’estimation statistique itérative ainsi que les progrès les plus récents, comme la modélisation du temps-de-vol des photons, ou la modélisation de la réponse du système dans le processus itératif.

Ces améliorations, dans un cadre statistique robuste et simple, impliquent en particulier une préparation des données en amont. La production par le scanner d’estimations correctes des projections de l’activité réelle est une condition pour la consistance de la reconstruction et des modèles de dégradation de l’image. Le mode d’acquisition passif est encore peu performant par rapport à la quantité d’information produite par la décroissance, et l’intégration nécessaire massive est limitée par des propriétés de saturation. Une partie importante de cette thèse se consacre donc à l’étude en amont de l’instrumentation de détection en utilisant des critères globaux de performance, définis par des standards internationaux.

Nous utilisons pour cela une série unique de mesures sur cinq modèles de caméras TEP commerciales, testées sur leur lieu d’utilisation clinique au cours de quatre années. Les analyses de résolution spatiale, temporelles et en énergie montrent la stabilité de cette technologie pour la représentation précise de l’information. Des analyses de sensibilité montrent un gain important au cours des générations étudiées ici et les progrès de l’électronique d’acquisition. Enfin le problème particulier de la performance en termes de restauration de contraste sur l’image TEP est discuté.

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I Un aspect important de la méta-utilisation de données de performance est la normalisation des résultats pour une comparaison correcte. Plusieurs méthodes sont proposées et appliquées en utilisant des propriétés temporelles des taux de comptage, qui sont en général basées sur les propriétés asymptotiques de la détection électronique de coïncidences fortuites. Ces trois familles de critères ne prennent que partiellement en compte un aspect moderne de l’imagerie clinique, qui est l’utilisation de l’information de différence de temps-de-vol des photons d’annihilation. Une mesure intrinsèque de résolution temporelle n’est que peu représentative de l’information au sein du processus de détection utilisée pour l’image clinique. Partant de ce constat nous proposons une technique de caractérisation inspirée des autres critères de performances et se basant sur les projections acquises pour une source ponctuelle. En testant la discrimination temporelle de l’information et le triage de l’information réalisé par les appareils et posant les bases du modèle de temps de vol en 2D, nous montrons sur un modèle de scanner TEP récent que plus de 40% des évènements ne suivent pas le modèle analytique idéal correspondant à une résolution temporelle idéale. À l’instar des mesures de fractions de diffusé, cette valeur fournit une information réaliste de l’impact de la technologie.

Ces deux études sont principalement focalisées sur la caractérisation des performances absolues des appareils. Nous ramenons ensuite cette limite en termes de performance de résolution de l’appareil à l’imagerie médicale, et en particulier la représentation de petites lésions. Pour ces objets, la résolution limitée de l’appareil et surtout les prérequis pour l’échantillonnage de l’image causent une sous-estimation importante de l’activité. En utilisant un modèle de la résolution intrinsèque de l’appareil il est possible d’utiliser des techniques de restauration itérative de la quantité d’information dans des lésions actives. Nous proposons une approche originale basée sur le traitement d’un modèle analytique d’une petite lésion et d’un modèle réaliste de la résolution spatiale de l’appareil. Les paramètres de cet algorithme peuvent être déterminés par les méthodes de mesures de performances détaillées précédemment. Nous étudions le schéma général de cette technique, peu sujette au bruit et à l’échantillonnage de l’image car opérant dans l’espace continu.

Ses limites sont rapportées également, et en particulier la simplicité du modèle ellipsoïdal inhérent qui limite le champ d’application de cette méthode. La généralisation des problèmes de quantification pour des lésions de taille plus importante dans des situations réalistes est soumise à la problématique du mouvement. Le déplacement des lésions cause un flou de déplacement dû à la faible rapidité de l’intégration des évènements par le système de détection. En conséquence la quantification classique en clinique basée sur le maximum d’intensité de la lésion est biaisée par un effet fortement anisotrope de dégradation de l’image. En utilisant une technique commerciale de détection du mouvement et un fantôme dynamique original basé sur un tissu pulmonaire réel d’un porc d’abattoir, nous étudions la complexité du mouvement de multiples lésions artificielles dans le poumon. En particulier l’impact de l’atténuation basée sur des images CT est variable au cours du mouvement respiratoire. Plusieurs techniques classiques sont évaluées telles qu’une phase corrélée entre des images synchronisées TEP/CT ou l’image moyenne CT et nous rapportons l’effet sur la quantification des objets simulés. La tendance attendue est que le mouvement physiologique impose une stratégie particulière pour les lésions périphériques. En amont, nous discutons également une approche plus juste de la génération des données quantitatives TEP. En utilisant cette fois-ci un fantôme dynamique simple et un milieu atténuant neutre, nous démontrons une quantification plus exacte de lésions simulées basée sur un seuillage en amplitude du signal physiologique. Cette nouvelle approche permet un traitement optimisé du mouvement et une quantification proche d’une situation de référence sans mouvements. Des avantages importants sont observés sur le temps d’acquisition des images thoraciques et la qualité de l’image.

Ce travail de thèse est donc situé à la transition entre des fondements théorique du traitement du signal expérimental et l’imagerie diagnostique et justifie sa classification multidisciplinaire. Notre approche est moderne et se base sur des technologies récentes pour détailler plusieurs aspects de la chaine de l’imagerie TEP. La précision de la quantification et le critère de performance clinique sont utilisés comme indicateurs de qualité, avec des mesures objectives se rapprochant au plus des situations réelles. Ce travail est également la base de développements futurs, pour améliorer l’étude des performances d’appareils TEP, et d’anticiper les comparaisons pratiques de ces appareils par le physicien médical. Les propositions faites peuvent être la base de nouveaux standards de comparaison pour des technologies, basées sur des méthodes ou des protocoles expérimentaux originaux.

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Abstract

Performance characterization and development of quantitative procedures for PET-CT scanners PhD thesis, University of Geneva, Switzerland

by Frédéric Schoenahl

Positron Emission Tomography (PET) is an imaging technique for the functional exploration of humans invented in the early 1970s. It underwent significant developments in the latest 20 years.

The acquisition mode for a PET tomography device is based on the particular physics of positrons and detects the results of radioactive positron emitters decays. The annihilation photon pair resulting from an interaction of the positron with the environment has a typical spatial correlation which can be used by the photon detecting camera to reconstruct the original location of the disintegration. The medical application therefore consists in developing molecules marked with a positron emitter and understanding their distribution within the human body, and its spatial representation using computer equipment. It is today a recognized technique especially for tumor search using the glucose analog ¹⁸F-fluoro-deoxy-glucose (FDG). The study and development requires therefore a large field of competence, which is driven by criterions for both medical and functional research applications.

In this work, we recall modern developments of this imaging technique by following both the signal processing chain and modelization aspects for the reconstruction necessary to characterize the radiotracer in-vivo distribution. We detail in particular the developments allowing three dimensional PET imaging, which was a major driver for the development and clinical efficacy of the technique in the late 20^th century. 3D acquisitions with limited collimation for photon-detection are challenging, but provide significant advantages in terms of sensitivity thanks to the newest signal processing possibilities and advanced developments in modeling the physics. The spatial correlation of annihilation photons ideally allows a sequential processing of the detected events, and the quantity of radiotracer can be indirectly measured. This important advantage of the counting technique requires corrections at different levels; several of them are discussed in detail in this work. The event rates in PET is significantly lower than those in other modalities like CT, and the management of noise due to e.g. attenuation and scattering is a complex statistical problem requiring accurate modeling. The first part of this thesis therefore reports fundamentals for the 3D radiotracer estimation based on the acquisition inverse problem, and discuss current statistical estimation methods, including most recent technologies like time-of-flight difference incorporation, or inclusion of system response models in iterative reconstruction problems. This part discusses as well the necessary processing steps to compensate for partial availability of the information due to the limitation in scanner geometry or hardware limitation as e.g. count rate saturation.

The experimental part details a meta-analysis of performance measurements on commercial PET systems, based on the international recognized NEMA standards for three generations and five models of commercial PET cameras. The spatial, temporal and energy resolution analyses show the stability and progresses in performance over generations and models. Sensitivity measures characterize some important differences over systems with different designs and reproducibility of these values. Image based performance analysis for contrasts restoration is reported with typical NEMA phantoms. We derive an important aspect for consistent analysis of such results with the requirement of normalization. Several methods are proposed and applied by observing properties of the random coincidences count rate or intrinsic radioactivity of the used crystal in these models and a strategy is devised for valid comparisons. The studied figures show limitation as well for a practical interpretation in clinical imaging, a typical example is on the relevance of timing resolution figures. As a consequence we propose a new performance figure for time-of-flight management capable scanners. Based on a 2D analytic representation of time-of-flight differences modeling for a point source, the technique attempts to characterize the discrimination power of the data processing chain. We show on a state-of-the-art commercial scanner that up to 40% of events could be wrongly assigned based on time measurement, what represents a more realistic impact of the technology.

These two studies are mostly reporting the absolute performance of devices. Another part of that work reports in more details the spatial resolution properties of the camera and evaluates its impact on the representation of small lesions. For these objects, the limited spatial resolution of the

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III camera and the inherent sampling issues cause an important underestimation of their activity. By means of an intrinsic resolution measurement inspired from the above methods, we propose an original iterative restoration technique based on analytic modeling. We study the behavior of this method, called model-based partial volume effect correction (MBPVEC). The technique appears less sensitive to noise and image sampling by working in a continuous space. Its limitations are reported as well, and in particular the simplicity of the inherent ellipsoidal analytic model which prevents its use in every possible situation. The generalization of quantification performance issues for small and larger lesions is discussed as well, based on similar analytical simulations. In a final part we discuss potential larger lesion image degradations caused by internal motions in a patient body. Due to the relatively slow integration capacity of the camera, lung movements impact the spatial representation of lesions with a motion blur. As a consequence the classical clinical image evaluation based on maximum intensity is subject to bias. Using an elaborated phantom design we first evaluate different the impact of varying attenuation properties due to motion of the tissues.

This first analysis is based on the conventional phase-based gating technique on PET data available today for commercial equipment. We used a phantom based on real porcine lungs, which can be animated with a realistic motion. Quantitative results show large variability and interestingly do not favor a general method for attenuation correction. We therefore experiment a new approach to PET gating to provide more precise quantitative results. The adaptive amplitude- based sorting of events based on a physiological signal provides a better framework for a diagnostic in PET imaging. This approach is simple, fast and less sensitive to respiratory changes.

A phantom studies shows at the well a quantitative performance comparable to a non-motion situation and significant change in patients cases management. This encourages more clinical validation of this method for a routine utilization in oncology scans, and promising other application fields.

An important criterion for image quality in this thesis was the accuracy of the quantification with objective measurements as close as possible to the clinical practice. The experimental parts provide performance characterization and a unique basis for further developments; it allows as well preparing and evaluating future and existing PET equipment by medical physicists. Further proposals and new methods in this work prepare future developments and software approaches for better data management and processing. This dissertation devises the transition between instrumentation, physics and the clinical practice of a medical imaging modality, justifying its classification as a multidisciplinary dissertation.

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I List of abbreviations

AAG Adaptive Amplitude Gating

BGO Bismuth Germanate (Bi2Ge3O12:Ce)

CT Computer Assisted Tomography (modality)

CTAC CT based attenuation correction

DIFT Direct Inverse Fourier Transform

DOF Distance-of-flight

DOI Depth-of-interaction

EM Expectation-maximization algorithm

FBP Filtered Backward Projection

FORE Fourier Rebinning

FOV Field-of-view (transaxial if not mentioned)

FWHM Full-width at half-maximum, referred to as w

GE Healthcare General Electric Healthcare (Company)

GEM Generalized EM

GSO Gadolinium Orthosilicate (Gd2SiO5:Ce)

GTM General Transfer Matrix

HPD Hybrid Photo Detectors

LABR3 Latrium Bromide (LaBr3:Ce)

LOR Line of Response

IGRT Image guided radiotherapy

LSO, LYSO Lutetium Orthosilicate (Lu2SiO5:Ce), doped (Lu0,6Y1,4SiO0,5:Ce)

ML Maximum likelihood

MRI Magnetic Resonance Imaging (modality)

MSRB Multi-slices Rebinning

MBPVEC Model based PVE Correction

MRAC MR-based Attenuation Correction

NECR Noise-Equivalent Count-Rate

NEMA US National Electrical Manufacturers Association

OSEM Ordered Subsets EM

PET Positron Emission Tomography (modality)

PSF Point spread function

PVC, PVE Partial Volume Correction, Effect

ROI Region-of-interest

RPM Respiratory cycles Per Minute

SPECT Single-Photon Emission Computed Tomography (modality)

SSRB Single-slice-rebinning

TOF Time-of-flight

TTD True-to-delayed coincidence rates ratio

VOI Volume-of-interest

Discovery®, Gemini® and Biograph® are commercial product lines of PET-CT scanners models.

II List of symbols

A, A^T,A^-1 The matrix A, its transpose A^T and inverse A^-1

u, u^T N-dimensional vector u, and its transpose u^T

X A random variable X

( ) A common function g

( ) A transform of g

i^(t) Instance of vector i at iteration t

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VII

III List of figures

Figure 1. Data flow and coincidence circuitry model.. ... 2

Figure 2. Acquisition coordinate system conventions for a PET system. ... 6

Figure 3. Data truncation problem when clustering angles for 3D acquisitions. ... 7

Figure 4. Acquisition systems based on projections and representation of data in the cylindrical space ... 9

Figure 5. 2D reconstruction problem for common geometries. ... 10

Figure 6. 2D coordinate systems and transformations showing reconstruction steps for a single projection... 11

Figure 7. Reconstruction of image quality phantom data without filtering term ... 13

Figure 8. High pass filtering for discrete back-projection reconstruction using a cylindrical phantom. ... 14

Figure 9. 3D approximate discrete rebinning techniques for PET ... 16

Figure 10. 3D coordinate systems and transformations. ... 16

Figure 11. Orlov representation of the -sphere and for a typical scanner acquisition geometry ... 18

Figure 12. Overview of available rebinning techniques using an IEC/NEMA image quality phantom ... 20

Figure 13. Quadrant system for the use of symmetries in 3D reconstruction ... 21

Figure 14. Ray-tracing models for forward or backward projection models in for the system model ... 24

Figure 15.Illustration of the OSEM convergence process ... 27

Figure 16. Sources of quantification errors from various parts of the image production chain ... 32

Figure 17. PET system response model for objects of homogeneous intensity ... 35

Figure 18. Random coincidence determined by delayed coincidence windows. ... 38

Figure 19. Conventions for plane-wise random coincidence rate variance reduction ... 39

Figure 20. Decay equation explaining the production of + radiations ... 40

Figure 21. Impact of magnetic field on the positron emission density distribution using simulations ... 41

Figure 22. Monte-Carlo simulation of annihilation photons in a 3D geometry ... 42

Figure 23. The arc- or geometric correction problem. ... 43

Figure 24. Typical quantitative problems for 3D PET acquisitions. ... 44

Figure 25. The attenuation problem. ... 45

Figure 26. Linear attenuation coefficients in cm^-1 for H2O ... 46

Figure 27. Phantom for the HU-to-511 calibration and attenuation correction example ... 48

Figure 28. µ-map derived from CT using dual energy mapping technique ... 49

Figure 29. Dominant interaction cross-section for different tissue effective atomic numbers ... 52

Figure 30. In-detector Monte-Carlo scattering modeling for a 1093 TruePoint design using LSO blocks ... 54

Figure 31. Monte-Carlo simulations of photon energy spectra for point sources of ¹⁸F and ⁸⁶Y. ... 55

Figure 32. Scatter contribution estimates in each polar representation using a simulation technique ... 58

Figure 33. Scatter correction in the FBP and OP-EM reconstructions for data of a TruePoint scanner ... 59

Figure 34. Brain HPD-PET concept ... 60

Figure 35. Energy profiles for the 8 first crystals of a LSO block ... 64

Figure 36. NECR analysis base data for a single frame ... 67

Figure 37. Possible normalization criterions for NECR curves comparisons ... 68

Figure 38. Timeline for accurate IQ phantom preparation using two identical preparations ... 70

Figure 39. NEMA 2007 Activities at start as an indicator for repeatability of preparations ... 71

Figure 40. Typical count-rates repeatability indicators over contrasts for two list-modes of 30min each ... 72

Figure 41. Decays and notations for two shifted acquisitions and impact of the isotope half-life ... 74

Figure 42. Ill-conditioned measurement with two series and no possible overlapping of series TTDs ... 75

Figure 43. Block-energy resolution dispersion ... 79

Figure 44. Dispersion and median statistic of energy resolution distributions ... 79

Figure 45. Standard error to the mean for energy resolution measurements ... 80

Figure 46. Block-timing resolution dispersion for 5 scanners of type 1103 and 1104. ... 81

Figure 47. Spatial resolution average values... 82

Figure 48. Paired design results for similarity testing of two different voxel sizes reconstructions. ... 83

Figure 49. Normalized NECR average and standard deviations for 1103 and 1104 models ... 86

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Figure 50. Normalized scatter fraction average trend and standard deviations for 110x models. ... 87

Figure 51. Hot sphere recoveries (sizes of 1.0 to 2.2 cm) ... 88

Figure 52. Relative noise within tumor to background level for NEMA spheres (1.0-2.2 cm) ... 88

Figure 53. Variability within the lung equivalent material around the central slice for both contrasts ... 89

Figure 54. Recovery coefficients and relative noise to background for the cold spheres (2.8, 3.7 cm) ... 90

Figure 55. 8:1 Recovery coefficients vs. standard deviation for all hot and cold spheres ... 91

Figure 56. 4:1 Recovery coefficients vs. standard deviation for all hot and cold spheres ... 92

Figure 57. Sample acquisition and calculation of efficiencies to real activity ... 93

Figure 58. Accuracy of point source positioning for a single operator ... 94

Figure 59. SSRB profile illustrating a filling error (air gap). ... 96

Figure 60. Representation of the DOF for two photons ... 102

Figure 61. DOF for an Y-offset source at b=10, 20, 30, 40 cm and verification of the analytic model ... 103

Figure 62. Verification of the analytical model using analytic simulations ... 104

Figure 63. Time and energy resolution stability for a 1104 scanner (typical) using a line source ... 106

Figure 64. DOF model segmented calculated for a source at 30 cm ... 107

Figure 65. Sinograms and time segments for a point source at (0,20) cm. ... 108

Figure 66. Point source scattering over segments for a point source source ... 109

Figure 67. Corrected assigned counts percentage ... 110

Figure 68. MBPVC algorithm description and the 10-parameters estimation problem ... 113

Figure 69. Acceleration via cumulated Gaussian functions ... 115

Figure 70. 3D visualization of the estimate from the MBPVC algorithm ... 116

Figure 71. Analytic simulation with Gaussian noise and a FWHM of 3 voxels ... 116

Figure 72. PSF model parameters measurement for the ECAT ART ... 117

Figure 73. inverse Quadratic penalty function chosen for the optimization process on real data ... 119

Figure 74. Convergence for a simulated lesion with selected background and recovery ratio ... 120

Figure 75. Study of the algorithm convergence on reconstructed patient data ... 122

Figure 76. 3D Activity-to-background recovery model for a spherical lesion of 10 mm ... 123

Figure 77. Spill out models for 2 mm, 6 mm and 10 mm FWHM Gauss-PSF filters ... 124

Figure 78. Theoretical recovery curves for difference sphere volumes ... 125

Figure 79. ArtiChest ® prototype for the realization of the gating evaluation study ... 129

Figure 80. Maximum intensity projection of 8 tumors of acquisition 5 used for analysis. ... 131

Figure 81. Adaptive amplitude gating principles ... 132

Figure 82. Motion simulation design for amplitude gating testing ... 132

Figure 83. Displacement of lesions over 8 gates for lesions of Figure 80 ... 134

Figure 84. Quantification performance over gates ... 134

Figure 85. Significant correlations between relative displacements from gate 1 for 12 lesions ... 135

Figure 86. Quantitative results over gates for the alternative 12 lesions of Figure 89 ... 136

Figure 87. Optimum search for quantitative performance ... 137

Figure 88. Average –CT AC compared to the absolute lesion location for a group of Figure 80 ... 138

Figure 89. Z-projection of 14 lesions used for static mismatch evaluation ... 138

Figure 90. Simulated mismatches using series of static acquisitions ... 140

Figure 91. Transaxial slices of Average CTs in the diaphragm region... 141

Figure 92. Qualitative impact on images with two different contrasts simulated in air balloons ... 142

Figure 93. Amplitude histograms of a patient respiratory signal and data utilization ... 143

Figure 94. Diagnostic performance in terms of quantification and detection using select patient examples .. 145

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IX

IV List of tables

Table 1. Main properties of some commercial and experimental scintillators ... 5

Table 2. Scanners recruited for evaluation of low-activity absolute sensitivity meta-study ... 65

Table 3. Settings for recovery calculations acc. NEMA 2007. ... 76

Table 4. Characteristics of the 38 Siemens PET scanner types recruited for the large scale analysis study. ... 78

Table 5. Constancy assessment for spatial resolution measurements on TruePoint systems ... 81

Table 6. Constancy assessment for spatial resolution measurements on mCT systems ... 82

Table 7. Statistics for absolute sensitivity measurements ... 84

Table 8. Noise equivalent Count rates and scatter mean statistics ... 85

Table 9. Axial source filling quality evaluation by voluntary addition of residual drops in the capillary ... 95

Table 10. Summary of significances for a difference in resolution figures between 109x and 110x ... 97

Table 11. Average gain matrix for absolute sensitivity between generation of cylindrical scanners ... 98

Table 12. Average gain matrix for clinical sensitivity (NECR performance) between generations. ... 98

Table 13. Measured sources and relative amount of wrongly assigned counts due to TOF performance ... 110

Table 14. Sphere volumes for the test objects used in the analytic resolution modelling study. ... 117

Table 15. Benchmarks between classes of numerical integration algorithms for MBPVEC. ... 118

Table 16. Recovery results for a constant FWHM. ... 120

Table 17. Summary of translational and scaling parameters restoration ... 121

Table 18. Absolute variations as average of all lesions relative quantitation variation for gating evaluation.139 Table 19. Max activity values for three attenuation correction techniques for the porcine lungs model ... 141

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1. Introduction 1

1 Introduction

1.1 Historical Perspective

Positron Emission Tomography (PET) (Hoffman and Phelps, 1979) is a recognized imaging modality for the analysis of functional processes within the human body. The practice of PET is actually inherited from the major modality used in nuclear medicine: Single-Photon Emission Computed Tomography (SPECT). These instruments detect gamma photons arising from electron capture process of chosen isotopes, free or coupled to a pharmaceutical. The high specificity of the molecule or the isotope itself for a certain organ, or metabolic process will label its target and locally emit photons, which can then be detected to attempt to localize the origin of emission. PET has had a long development towards clinical use, and its wide use today has been preceded by SPECT, which is simpler to set-up despite a lower intrinsic performance. The initial development of PET mostly occurred on the western and eastern coasts of the United States at university centers. Initially the technique was a functional research tool and required an isotope production and research facility at a close location. The impacting reported applications were about the physiology of animal models and the study of major biological functions, in particular the brain and the heart. However PET slowly adopted the appearing clinical concept of nuclear medicine imaging, and some attempts were made to perform scintigraphic examinations on very basic equipment with opposed detectors equipped with coincidence electronics. For both instrumentation and clinical use, PET has been coupled to the development of SPECT, and has followed the way to a clinical specialty SPECT pioneered, to become equally considered today among nuclear imaging techniques. Interestingly both modalities have evolved very slowly until today, and very gradually. Their cost is still among the highest for imaging and the logistic an important issue. Nuclear medicine is probably the domain which has linked the latest the medical constraints to its use. It is useful to not separate PET from SPECT when discussing the historical dimension of this imaging modality. In fact, without knowledge of the medical problem underlying a specific nuclear image, its assignment to one of the two modalities is an impossible task. The qualitative evaluation of nuclear images is not providing the relevant information.

The role of instrumentation, since its first use, was to maximize the confidence in detectability. The specificity of the functional process is inherent to the tracer, and cannot be solved by instrumentation itself, or using unstable inference processes. Detection of the tracer within the field-of-view (FOV) of a SPECT or a PET camera is actually what guided more than 30 years of instrumentation development. Above the concept similarities between the two techniques, there is a remarkable difference. The underlying PET physics allows a fundamental spatial and temporal correlation of the acquired information. This was discovered early as being a major advantage over SPECT as this correlation virtually allows quantification of the information present within the FOV. Ideally PET is a real spatial and temporal counting device which allows producing quantitative and dynamic images. In the last decade the use of PET has spread as fast as SPECT, which is remarkable regarding the significantly higher costs for a PET installation. This somehow demonstrates as well the importance of this quantitative aspect. And this is further supported by the current researches on SPECT cameras for developing quantitative techniques.

1.2 Multimodality imaging: more than the sum of its components

When performing SPECT imaging, the photon birth place is estimated by the mean of collimation techniques, to produce a planar projection of the current radioactivity distribution. Collimators would absorb a random amount of incoming photons with non-normal incidence angles. The two techniques, PET and SPECT, differ essentially with the pattern of radiotracers which can be chosen to target different physiological function and the level of complexity of the inherent instrumentation performance needed (Budinger, 1992, Rahmim and Zaidi, 2008). Positron based nuclear imaging can both reuse collimation techniques (2D mode) or purely exploit the geometric properties of positron annihilation which results in the collinear emission of photons in opposite directions. Both can only be achieved with detectors located in opposition around the patient and ideally surround completely the imaged subject. The accumulation of connected detections or so-called lines-of-responses (LOR) allows an estimation procedure to guess the location of the positron along this line in a 3D space.

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This strong correlation enables accurate localization at the costs of a much more demanding hardware performance to allow testing and connecting a massive amount of local detections.

Both nuclear imaging techniques are based on photon acquisition and therefore use the same concept of imaging chain. The main component is the scintillator crystal, included in a so-called detector block. The crystal material basically converts photons into visible light, carried into light guides to later be amplified for the purpose of electronic signal generation. This is achieved via transformation of visible light in electrons. Electrons are then amplified in vacuum by successive collisions onto dynodes elements under a chosen voltage potential, or alternatively using an avalanche effect in an amplification semi-conductor block. Electrons are then further carried to an integration circuitry and complex digital block capable of comparing channels with time synchronization, and prepare the logging of information into physical data streams. Figure 1 shows an overview of the instrumentation concept.

FIGURE 1.DATA FLOW AND COINCIDENCE CIRCUITRY MODEL, SHOWING THE NECESSARY PROCESSING FOR A SINGLE PAIR OF DETECTORS.THE REST ENERGY SIGNAL OF PHOTONS FOR EACH DETECTOR CHANNEL IS EVALUATED TO BELONG TO AN ENERGY WINDOW[ ].BOTH SIGNALS ARE THEN TIMELY CORRELATED BASED ON THE SCINTILLATION TIME, AND THEIR ARRIVAL TIME DIFFERENCE COMPARED ^{TO A}

DEFINED TIME WINDOW .

Based on coincidence information stream, the accumulation and integration of these events over time provides the material for the fundamental operation of radiotracer distribution estimation. This is the base for diagnostic which should display the labeled pharmaceutical in a Cartesian reference with the highest possible accuracy. This is as well where PET strongly differentiates from SPECT, as the spatial correlation mentioned above virtually allows the technique to relate the amount of radiation detected with the real amount of radiotracer uptake, where SPECT remains blind to all escaped events.

Radiotracers are the key for the success of all nuclear imaging modalities; they however are costly because of complicated production, transport and administration logistics. They provide the modality which a unique specificity to a functional process or disease, such as a tumor type at a particular stage of its development. The abundant availability of a large number of radiotracers is promised since the beginning of the 21^st century; however the development is extremely slow due to licensing issues which can be a decade work as a radiotracer is falling under pharmaceutical rules.

Coincidence energy

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1. Introduction 3 Commercial distribution networks or cyclotron activities from universities and private parties for the dispatching of isotope are still in development phase, especially outside the USA. Despite these difficulties for expansion, PET remains among the highest sensitive imaging modalities against classical radiological modalities as Computer Tomography (CT) or Magnetic Resonance Imaging (MRI). The sensitivity in terms of detectable concentration of tracer for PET significantly outperforms other modalities, and is estimated to 10^-10 – 10^-12 mol. It remains 10^-2 for CT assuming the use of contrast agent, and 10^-3-10^-5 mol (in proton rich areas) for MRI (Jones, 2002, Nestle et al., 2009). Even if these modalities have themselves potential for functional imaging too, the availability of a PET component in an hybrid anatomic-functional device like PET-CT (Beyer et al., 2000, Townsend, 2001) or PET-MR (Shao et al., 1997, Zaidi et al., 2007a, Pichler et al., 2008, Pichler et al., 2010) is a clear advantage for finer detection and quantification of disease, and outstanding research in major clinical fields or for drugs development.

1.3 Aims of this work

In this work we will recall the different aspects of instrumentation leading to the definition of criterions for performance evaluation, and develop methods for improving the quantitative accuracy in clinically challenging situations related to PET imaging. This thesis is built with two large sections. The first section of this work covers theoretical aspects of modern PET imaging, and reports in three parts the important developments which led to the routine practice of the modality. A second large section is divided into four distinct experimental parts, and reports analyses and developments related to this technology.

Within the first section, we first set the focus on discussing modern 3D PET instrumentation, and the theoretical requirements for its exploitation to its utmost potential. The theory of 3D image reconstruction is then extensively discussed, starting from the classical general description of an analytic solution (filtered backward projection, FBP) and towards current iterative estimation algorithms (generalized EM). The problem of image reconstruction involves sources of inaccuracies from various origins. The next part reports major factors resulting in image degradation, including instrumentation related aspects. A separate part is dedicated to the two classical photon interaction physics problems, attenuation and scattering, with a detailed review of the correction techniques and modeling aspects.

In the second section reporting experimental studies, we first use a unique large panel of PET cameras and define acceptance ranges according to a widely accepted standard. The results are both hardware and image quality figures. We especially focus on the problem of normalizing results and propose original methods for comparing heterogeneous acquisition situations. We ensure therefore requirements are set for a valid comparison over and within systems. The second experiment is a proposal to extend the performance tests with a time-of-flight evaluation based on an image generation criterion and a scattering index. This novel technique is for a better characterization of the timing resolution on the first generation of commercial scanners equipped with time-of-flight based reconstruction. In a third experimental study, we propose to overcome the limitation of one of the hardware figures, spatial resolution, by using an iterative estimation method for the recovery of small lesions activity and shape parameters directly on images. The technique is evaluated using analytical simulations and some patient images. We further extend the PET image processing aspects in a last experimental part, with a focus on spatial resolution degradation due to subject dependent motion in the thorax region. A novel dynamic phantom is used to quantify precisely the impact of the attenuation correction for this classical situation based on clinical indices. The phantom based on real porcine lungs allows a realistic evaluation of patient thorax analyses. In order to simplify the motion management, we propose an alternative technique for motion freezing based on data sorting with the acquisition and segmentation of the physiological signal in amplitude. The quantification accuracy of the novel technique is demonstrated in comparison to the current clinical standard which uses phase selection.

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2 Overview of modern clinical PET-CT instrumentation

2.1 State-of-the-art PET-CT instrumentation

The common purpose of nuclear emission imaging and core technical difficulty is the three dimensional estimation of the distribution of a radioactive tracer as close as possible to a real distribution within an imaged subject. The imaging process is a rich workflow which can be divided into a data gathering part, or acquisition, and an estimation part, or reconstruction (Vandenberghe et al., 2001). Both are usually taken apart to distinguish research applications in instrumentation (optimization of data acquisition schemes) or image processing (optimization of the estimation problem). Both steps are conditioned by both the clinical question and the inherent physics of the emission and detection processes, and the estimation step integrates this knowledge to improve the estimation by means of accurate modeling of charged and non-charged particle-material interactions.

These are in particular photon interactions (elastic and non-elastic interactions) with both the patient and the detection device, positron interactions within the patient, and the physics of radioactive decay.

Instrumentation was historically focusing on addressing every component of the classical nuclear medicine acquisition chain as proposed early by Al Anger et al. (Anger, 1952, 1953). The scintillation crystal for photons detection, multiple photomultipliers, and digital back-end electronics for coincidence detection all underwent dramatic improvements which led to the realization of accurate functional investigations in reasonable time. In the history of instrumentation, the crystal always played a central role. The complex step of estimation is a statistical process based on little information summarized by the individual and sequential scintillations caused by single photon interactions within the crystal structure. We can identify different steps in the estimation process, first by identifying the impact of hardware limitations on the counts identification and selection, and then purely the software parts which shall solve an inverse problem and involve several corrections for both quantitative and qualitative aspects.

Commercially available scanners today use large number of cuboids arranged around a cylindrical field-of-view. The detector rings are usually designed to avoid gaps between the different scintillation crystals, such that a trade-off between radius and number of crystals provides a uniform coverage of a certain solid angle. Looking closer at the ordering of crystals, the detection surface is hence polygonal. Pixelization of the information using multiple facets is another difference with SPECT, where the localization of information is diffuse within a single large crystal and the pixel identification made by the analysis of the light dispersion (Anger logic). Such an approach can however improve the scintillation localization and avoid certain data assignment problems, like the joint project for transferring the Hybrid Photo Detectors (HPD) technology from the High Energy Physics Experiment at CERN to an optimized multilayer design for brain PET imaging (Braem et al., 2004).

2.1.1 2D mode imaging

The historical way to reduce the complexity of each scintillation measurement has been inspired by single-photon emission tomography methods, and consists in eliminating non-orthogonal photon incidences on the detectors using very dense shielding materials (typically lead or tungsten). The counting capacity in such devices is very limited as numerous incidences will be absorbed within the lead septa hence limiting the load on the acquisition electronics. In PET, 2D acquisition uses circular collimation rings and is followed by a plane wise reconstruction technique to generate slices as for classical sequential CT. The septa shall in fact be built between each detection pixel or crystal unit in the axial direction for being really efficient; however the complexity of the crystal matrix usually does not allow such a fine collimation, which remains purely axial on many classical models (Spinks et al., 1992). The overall absolute sensitivity of such geometry, defined as the ratio between injected dose and the number of counts integrated by the system is relatively low, thus requiring higher patient doses or longer scan times to ensure consistent statistics. This acquisition mode is; however still in use in several institutions and marketed in several products for the ability to provide less noisy images in specific situation such as thoracic and cardiac scans where scattering and/or random events add significant noise to the acquisition data. Added to a low sensitivity performance, another disadvantage is the anisotropic discrimination of coincidences. The shielding is not isotropic (like a pinhole concept in SPECT imaging) and attenuates axially oblique coincidences to hit the detection surfaces. If the shields are more likely to e.g. discriminate non-collinear photons of a pair due to the

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2. Overview of modern clinical PET-CT instrumentation 5 scattering of one or both of them, useful non-scattered oblique coincidences will be eliminated and the solid angle is extremely small, as it covers a single crystal in the collimated direction, typically axial. This mode has been available in various commercial models between 1990 and today. This mode is essentially further developed and sold by General Electric company (GE) for its BGO Discovery™ product line, with some variations towards reduced or partial collimation (Schmitz et al., 2007, Alessio et al., 2008).

2.1.2 Fully 3D mode

Computer innovations in the last decade have provided a dramatic increase in computational power, data integration speed and capacity. Combined with the better management of the imaging chain and a better knowledge of the detection physics, a fully 3D concept was proposed early based on an affordable crystal, BGO (Cho and Farukhi, 1977, Townsend et al., 1991, Badawi, 1997). The 3D acquisition slowly supersedes completely 2D modes, as it uses exponentially more of the solid angle of emission for photons. As a consequence the integrated information for an equivalent patient dose and acquisition duration is significantly higher. Early images have been criticized for being largely affected every kind of physical annoyance due to the physics of emission and detection.

The choice of the crystal became a bottleneck for the amount of data, which could have been integrated in a time unit. The ideal crystal to perform 3D imaging had to have a fast scintillation light decay, possibly a large amount of light output and a high stopping power to ensure the actual absorption in the crystal volume. A necessary constraint is the ability to manufacture and obtain stable performances, such that only a few candidates match more of these requirements. This mode has been impacted by the development of LSO (Nutt, 2002, Nassalski et al., 2007), which has been a proprietary development and exclusive to CTI Systems and Siemens Healthcare until 2008. The companies integrated efficiently the crystal with fast digital electronics around 2003 for 3D only acquisitions. Noticeably, two over three manufacturers (Philips Healthcare and Siemens Healthcare) do not supply a product with a 2D mode anymore; GE Healthcare’s research tomograph the Discovery™ 690 (January 2009) was the first for this manufacturer to join the family of full 3D scanners. All manufacturers use now a LSO crystal with some design variations; Philips Healthcare and GE Healthcare use an yttrium doped LSO (LYSO), with slightly different detection performances which non significantly differ from the original LSO for their latest realizations (respectively Gemini™ TF and Discovery™ 690). LYSO is manufactured and sold by St. Gobain Crystals, USA. Siemens Healthcare still continues the original LSO concept with a cerium doping to increase the crystal efficiency (Table 1).

LSO LYSO BGO GSO LaBr3

Light output [ph/MeV]

31000 32000 8500 7600 65000

Peak emission [nm]

420 420 480 430 360

Decay time [ns]

40-47 41 300 30-60 15

Refr. Index 1.82 1.81 2.15 1.85 1.9

Density [g/cm³]

7.4 7.1 7.13 6.71 5.29

TABLE 1.

MAIN PROPERTIES OF SOME COMMERCIAL AND EXPERIMENTAL SCINTILLATORS.LSO,LYSO AND BGO

ARE PRESENT WITHIN ALL RECENT MANUFACTURED PRODUCTS IN 2011.GSO-EQUIPPED SCANNERS ARE STILL PART OF THE CURRENT INSTALLED BASIS.LABR3 IS OF PARTICULAR INTEREST FOR TIME-OF-FLIGHT

TECHNOLOGY DEVELOPMENT, WITH FAST DECAY TIMES.FROM NASSALSKI ET AL.(1997).

The first problem appearing with the possibilities of performing fully 3D PET acquisitions was to manage the large amount of data generated by all possible coincidences integrated by the

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electronics. Both in terms of storage and reconstruction, the requirements in memory and disc space were not tractable with native data. In order to reduce the amount of storage space needed for 3D data, projections acquired during the PET examination needed to be compressed with minimal losses of information. A proposal by Dr. C. Michel et al from the Catholic University of Louvain-la- Neuve, Belgium was a systematic resampling of polar data in the three directions of a cylindrical space (Figure 2).

Based on the available sampling for a given tomograph, the compression method uses conventionally visualization along tangential directions (views in a sinogram), radial directions (projection bin in a sinogram), and axial directions (planes in a sinogram). In fact only tangential and axial data representations are potentially really compressed, as the data reduction radially is usually achieved by trimming projections in areas where data is irrelevant (edges of the transaxial field of view). The radial trimming will not be discussed in details, and is analog to a processing in a

(a) Cartesian/polar coordinate conventions. (Left) 3D view of the PET gantry bore, (middle) a transaxial view and (right) a sagittal view.

(b) Cartesian representation of the polar discrete projection space with view- and planagrams defined around the Z axis and sinogram defined normally to the isocenter z-axis.

FIGURE 2.ACQUISITION COORDINATE SYSTEM CONVENTIONS FOR A PET SYSTEM WITH DETECTOR RADIUS R_D

WITH (a) THE CARTESIAN AND POLAR CONVENTIONS (b) A COMPRESSED DISCRETE REPRESENTATION OF THE

3D PROJECTION SPACE IN THE FORM OF A 3D VOLUME P(R,Θ,).

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2. Overview of modern clinical PET-CT instrumentation 7 different context (later in this work, Figure 36). Left and right margins of projection sets are simply removed in order to reduce the dataset size, other compression operations would degrade significantly the image quality as they would affect directly the image sampling. The trimming of projections at a cutoff distance r_c from the isocenter¹, with b the bin size, can be expressed as

( ) | | ← ( ) | |

The dataset is reduced accordingly with a factor , where is the number of projections. This is a lossy technique, and this simple approach is chosen for specific clinical applications as e.g. a brain examination where a large part of the transaxial plane is not covered by the subject image.

Contrary, the tangential compression is based on an increase in the angular sampling size by integer numbers and a summation of successive projections around the FOV. Often referred to as views mashing, the data reduction is possible at the cost of the sampling accuracy for the  parameter. The information is not lost, but degraded by this process through averaging. Using a mash factor 2^m, projections are transformed as follows and processed projections are removed from the dataset:

( )← ∑ (  )

Mashing allows a reduction with a factor 2^-m. Both radial and tangential compressions are 2D problems only, and can be realized for a single plane normal to the z axis of the scanner.

The removal of physical collimation allowed an additional set of detector combinations, with the same  parameter and an axial obliqueness angle  (Figure 3).

FIGURE 3.DATA TRUNCATION WHEN CLUSTERING ANGLES WITH (left) THE 3D TRUNCATION PROBLEM FOR FINITE AXIAL GEOMETRIES.IDENTICAL SAMPLE STEPS CAN BE SUMMARIZED BY AN AVERAGE ANGLE ^̅. FOR

̅ THE ACTUAL CLUSTER SIZE REDUCE WHEN APPROACHING THE MAXIMAL ACCEPTANCE ANGLE OF THE SCANNER (right) REAL DATA CLUSTERING FOR A NEMAIEC PHANTOM ACQUISITION AND MARKED

BOUNDARIES ACCORDING TO EQ.(3).

Over N_r detector rings in z, possible ring combinations are possible when the scanner is operated in 3D mode (Bendriem and Townsend, 1998). The span is the axial compression factor s and will

1 We assume here that the projections are corrected for geometric arc-effects and the bin size of the projections is homogeneous, otherwise the truncation may be larger than wanted

z

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

 z

p

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combine successive axially oblique projection combinations, analog to what mashing does tangentially. This is the only compression specific to 3D scanning, as it combines coincidences acquired from detectors located on two different rings. The largest axial coverage allows a largest angle _max which is the maximum value to consider for compression. The actual axial angle for a LOR can be expressed with respect to the difference in axial coordinates between for the two physical rings ( ) and the ring detector diameter R_d for a given design. By first defining the unit vector n along the LOR, its normalized coordinates for a LOR recorded between two rings at axial coordinates z1 and z2 can be written as

√ ( ) ( )

* ( ) √ +

(1)

With the left term a normalization quotient. The axial angle  is defined as:



√

(2)

Here r refers to the distance of a LOR with the z-axis. The unit vector is trivial to define in the 2D transaxial plane with a radial and tangential component along the LOR. In 3D the last axis can be represented as a normal along the LOR row. r is considering fading regarding the significantly larger detector ring diameter. The minimal angle to consider in this case is the normal plane to the axis, marked ₀. Negative ring differences are spanning [ [ and positive ring differences] ].

A set of boundaries are then calculated to group close angles together and span the 2max angle range. The angular span is divided in an odd number of segments, with a central group of segments around  . Starting from a span factor s giving the number of angle bins to be clustered together, discrete boundaries u can be calculated for successive segments. For segment i the lower and upper boundary for Ni segments are given as follows:

( ) with Boundaries are calculated by sorting indices by absolute values

{ ^{| |} ( ) ^{| |}

and {

^{| |} for ( ) for

| | for

(3)

These rules, called after the first commercial 3D scanner prototypes ECAT have been devised by Dr.

C. Michel from the Catholic University of Louvain-La-Neuve, who devised a conceptual representation of the axial compression concept. These are still in use on all modern 3D capable scanners for data reduction today. The gain is non-linearly depending on the span value chosen.

These compression rules have been implemented in an in-house C library (NLab) to allow interactions with all possible compressed projection set types, or produce compressed projection sets from simulation data for instance.

2.1.3 The nuclear imaging instrumentation chain

The sequential nature of acquired data best describes a nuclear tomographic device as a counting device. The detection unit common to emission tomography devices has two major tasks: detect single photons at any point of the acquisition geometry and correlate spatially the information in order to later attempt to identify the origin of LORs within the field-of-view (Phelps et al., 1975b).

The crystal volume is usually designed as a single layer polygonal detection surface, which can be simplified to a cylindrical photon absorption annulus. PET spatially correlates photon pairs hitting the surface of the cylinder at any point, and uses a pixelated structure of small crystals. Focusing on the detection unit and a single photon, the so-called Anger chain processes the impact of a photon up to usable information for geometric processing (Figure 4). The photon absorption unit is a dense inorganic and doped crystal, which favors photoelectric interaction of incoming photons, and converts its energy into visible light by electron capture. In the context of counting rate performance,

Performance characterization and development of quantitative procedures for PET-CT scanners

Thesis

Reference

Performance characterization and development of quantitative procedures for PET-CT scanners

SCHOENAHL, Frédéric François

SCHOENAHL, Frédéric François. Performance characterization and development of quantitative procedures for PET-CT scanners . Thèse de doctorat : Univ. Genève, 2011, no.

Sc. 4320

URN : urn:nbn:ch:unige-167694

DOI : 10.13097/archive-ouverte/unige:16769

UNIVERSITÉ DE GENÈVE

Performance characterization and development of quantitative procedures for PET-CT scanners

THÈSE

Résumé

Abstract

Table of contents

I List of abbreviations

II List of symbols

III List of figures

IV List of tables

1 Introduction

1.1 Historical Perspective

1.2 Multimodality imaging: more than the sum of its components

1.3 Aims of this work

2 Overview of modern clinical PET-CT instrumentation

2.1 State-of-the-art PET-CT instrumentation