An introduction to molecular imaging in radiation oncology ... · An introduction to molecular imaging in radiation oncology: ... images a patient’s biology instead of anatomy alone

An introduction to molecular imaging in radiation oncology: A reportby the AAPM Working Group on Molecular Imaging in RadiationOncology (WGMIR)

Michael T. Munleya)

Department of Radiation Oncology, Wake Forest School of Medicine, Winston-Salem, North Carolina 27157

George C. Kagadisb)

Department of Medical Physics, School of Medicine, University of Patras, P.O. Box 13273,GR 26504 Rion, Greece

Kiaran P. McGeec)

Diagnostic Radiology, Mayo Clinic, Rochester, Minnesota 55905

Assen S. Kirovd)

Medical Physics, Memorial Sloan-Kettering Cancer Center, 1275 York Avenue, New York 10021

Sunyoung Jange)

Radiation Oncology, Princeton Radiation Oncology Center, Jamesburg, New Jersey 008831

Sasa Muticf)

Radiation Oncology Department, Mallinckrodt Inst of Radiology St. Louis, Missouri 63141

Robert Jerajg)

Department of Medical Physics, University of Wisconsin, Madison, Wisconsin 53705

Lei Xingh)

Radiation Oncology Department, Stanford University School of Medicine, Stanford, California 94305-5847

J. Daniel Bourlandi)

Department of Radiation Oncology, Wake Forest School of Medicine, Winston-Salem, North Carolina 27157

(Received 18 October 2012; revised 19 July 2013; accepted for publication 10 August 2013;published 30 September 2013)

Molecular imaging is the direct or indirect noninvasive monitoring and recording of the spatial andtemporal distribution of in vivo molecular, genetic, and/or cellular processes for biochemical, bio-logical, diagnostic, or therapeutic applications. Molecular images that indicate the presence of ma-lignancy can be acquired using optical, ultrasonic, radiologic, radionuclide, and magnetic resonancetechniques. For the radiation oncology physicist in particular, these methods and their roles in molec-ular imaging of oncologic processes are reviewed with respect to their physical bases and imagingcharacteristics, including signal intensity, spatial scale, and spatial resolution. Relevant molecularterminology is defined as an educational assist. Current and future clinical applications in oncologicdiagnosis and treatment are discussed. National initiatives for the development of basic science andclinical molecular imaging techniques and expertise are reviewed, illustrating research opportunitiesin as well as the importance of this growing field. © 2013 American Association of Physicists inMedicine. [http://dx.doi.org/10.1118/1.4819818]

Key words: molecular imaging, functional imaging, oncologic imaging, radiation therapy

1. INTRODUCTION

To address educational and clinical issues for molecularimaging in radiation oncology, the American Association ofPhysicists in Medicine (AAPM) formed the Working Groupon Molecular Imaging in Radiation Oncology (WGMIR).This report from the WGMIR is for educational and ref-erence purposes, and reviews the terminology, basic scien-tific approaches, clinical applications, and national initiativesin molecular imaging as relevant to cancer diagnosis andtreatment.

2. TERMINOLOGY, NOMENCLATURE,AND DEFINITIONS

Molecular imaging is the science whereby cells, or bio-logical species such as proteins and molecules, are imagedor used to noninvasively image in vivo biological structureand/or function at the molecular or cellular level. Defini-tions of molecular imaging show varied scientific perspec-tives with common basic principles. In April 2005, the So-ciety of Nuclear Medicine [SNM, now named the Society ofNuclear Medicine and Molecular Imaging (SNMMI)] and the

101501-1 Med. Phys. 40 (10), October 2013 © 2013 Am. Assoc. Phys. Med. 101501-10094-2405/2013/40(10)/101501/23/$30.00

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101501-2 Munley et al.: Molecular imaging in radiation oncology: AAPM WGMIR report 101501-2

Radiological Society of North America (RSNA) held a sum-mit meeting on molecular imaging that included representa-tives of the AAPM and other scientific societies. From thismeeting a consensus statement was crafted to define “Molec-ular Imaging” as1

“Molecular Imaging techniques directly or indirectly monitorand record the spatiotemporal distribution of molecular or cel-lular processes for biochemical, biologic, diagnostic, or thera-peutic applications.”

A more descriptive definition was presented by the (for-mer) SNM alone.2

“Molecular imaging is the visualization, characterization, andmeasurement of biological processes at the molecular and cel-lular levels in humans and other living systems. To elabo-rate, molecular imaging typically includes 2- or 3-dimensionalimaging as well as quantification over time. The tech-niques used include radiotracer imaging/nuclear medicine, MRimaging, MR spectroscopy, optical imaging, ultrasound, andothers.”

Depending on the molecular imaging technique, the de-tected signal may arise from an endogenous material or anadministered material such as a contrast agent. Additionally,chemical or biological amplification of the number of signal-generating molecular species present may be manipulated toincrease signal intensity, and thus detectability. Clinical imag-ing modalities are considered molecular imaging when phys-iological, chemical, or functional imaging is performed thatimages a patient’s biology instead of anatomy alone. For in-stance, the physiological distribution of a radiopharmaceuti-cal in nuclear medicine imaging has a molecular basis, andthe signal intensity in magnetic resonance imaging, thoughof nuclear origin, is regulated by the local molecular envi-ronment and thus also has a molecular basis. Both of theseexamples conform to the quoted definitions of molecularimaging.

For radiation oncology purposes, molecular imaging thatimproves detection of tumor cells, including their spatial ex-tent, phenotype, and particular biological character, may beused to more accurately define targets and dose regimensfor individual patients. One molecular approach to radia-tion oncology has been described as a “three-block process”comprising molecular imaging, molecular signature deter-mination (e.g., molecular characterization of a tumor), andmolecular therapeutics.3 Molecular therapeutics in radiationoncology refers to the use of ionizing radiation to initiatespecific molecular events that are directly therapeutic them-selves or to manipulate pathways that lead to an intendedtherapeutic effect.3 Molecular imaging methods are also be-ing investigated as a direct monitor of the radiation deliveryaccuracy.4–7

Molecular imaging to assess radiation treatment responsemay provide the basis for prediction of tumor and normal-tissue responses, initially for populations of patients and laterfor individual patients as specific molecular signatures be-come available for clinical use. Tumor and normal-tissue re-sponse data may help in the formulation of biophysical re-

sponse models that could be used to optimize computer-basedtreatment planning.8

Clinical imaging devices and tools are commercially avail-able to acquire molecular images. For example, positron emis-sion tomography (PET) and magnetic resonance imaging(MRI) systems can provide image data sets that can be corre-lated with computed tomography (CT) for radiation treatmentplanning and verification. The role of molecular imaging incancer treatment will thus continue to increase.

Clinical imaging of a cancer target with conventional MRIand/or CT will show normal anatomy and the target regionwith variable enhancement due to natural tissue contrast orthe administration of a contrast-enhancing agent. The incor-poration of molecular, physiological, or other biological char-acteristics of a tumor into a treatment plan remains the respon-sibility of the clinician through clinical judgment, a complexprocess that assimilates laboratory and pathological reports,knowledge of spread of disease pathways, and expected clini-cal course. However, biological tumor characteristics, such asthe presence of hypoxia and proliferation rate, are importantdeterminants in radiation treatment response.9–11 Hypoxia in-dicates possible regions refractory to radiation treatment, andcell proliferation may indicate regions of active tumor growth.Thus, in vivo molecular imaging of cell hypoxia, proliferation,and other important tumor characteristics and properties is ofgreat interest to the radiation oncology community12–15 andwill enable diagnosis, radiation treatment, and evaluation inthe molecular manner described by Coleman and others.3, 16, 17

Of particular interest to the radiation oncology community isthe concept of molecular target credentialing where a specificmolecular target is imaged, its molecular signature defined, atreatment given, and the effect of the intervention and the sig-nature is then evaluated. This approach can be used to validatethe proposed target as a legitimate one for cancer therapy andto provide the opportunity for individualized therapy on thebasis of both the initial characteristics of the tumor and thetumor’s response to an intervention.3

3. MOLECULAR IMAGING MODALITIESAND TECHNIQUES

The molecular imaging field includes multidisciplinaryscientific expertise and multimodality imaging to produceanatomical, physiological, or functional images of biologicalprocesses over a range of spatial scales from molecular to en-tire organ levels. The number or density of molecular signalsdepends on the imaging technique and the targeted biolog-ical process or system. Definitions for the multidisciplinaryterminology relevant for this field can be found in Cox andNelson18 and Wagenaar et al.19 and specifically for oncologyin Pomper and Gelovani.20

Five modalities are typically used for molecular imag-ing: (a) PET, (b) SPECT, (c) MRI, (d) optical, and (e) ul-trasound. Table I lists these modalities and their correspond-ing contrast agents, and summarizes their salient physicalproperties and their respective advantages and disadvantages.These five imaging modalities and their techniques are nowreviewed.

Medical Physics, Vol. 40, No. 10, October 2013


TABLE I. Various modalities and contrast agents used for molecular imaging (see Refs. 238–244).

Spatial resolutionTracers and

Technique probes Signal detection Sensitivitya Preclinical studies Clinical studies Temporal resolution Advantages Disadvantages

PET Radiolabeledtracers

High-energyγ -rays (511 keVphotons)

10−15 1–2 mm 1.8–2.3 mm Minutes Very highsensitivity forfunctional,metabolic andmolecularprocesses

Poor spatialresolution,expensiveequipment,requires cyclotron,can monitor onlyone isotope,exposure toradioactivity

SPECT Radiolabeledtracers

Low-energyγ -rays

10−14 1–2 mm 3–5 mm Minutes High sensitivity,can detectmultipleradiotracers andimage multipleprocesses

Poor spatialresolution,expensiveequipment,exposure toradioactivity

MRI Gadoliniumagents(±radiotracers)

Radiofrequencywaves

10−9−10−6 50–100 μm 0.25–0.5 mm Seconds–hours High soft-tissueresolution andcontrast to noiseratio, dynamicstudies

Expensiveequipment

Optical Bioluminescentand fluorescentprobes

Light (nearinfrared)

10−12 1–10 mm . . . Seconds–minutes Highly sensitivefor molecularfunctions, noradiationexposure,inexpensive

Depth-dependentattenuation ofemitted light

Ultrasound Microbubbles(±antigen labels)

Ultrasonic waves 10−8 50 μm 0.5–1.5 mm Subsecond–minutes No exposure toradiation, quick,inexpensive

Poor imaging inpresence of air orbone

aSensitivity refers to moles of label detected.

3.A. Positron emission tomography

The basic physical event in the formation of PET images isa simultaneous detection of two photons produced by the an-nihilation of a positron within a targeted tissue in the patient.21

The positron-emitting radionuclides used are proton-rich iso-topes, mostly cyclotron-produced, with half-lives betweenabout 1 min and several days.22, 23 These radionuclides areattached to a biological tracer molecule designed to localizein vivo in tissues with specific properties such as increasedmetabolism (e.g., glycolysis, cell proliferation, and protein ormembrane synthesis), tumor hypoxia, or disease-specific tar-geting, as discussed in more detail below.23, 24

Positrons are emitted with a continuous energy spectrumwith a characteristic maximum energy from 0.63 MeV (for18F) to ∼ 3.7 MeV (for 76Br) among clinically used positron-emitting radionuclides. An emitted positron subsequentlyundergoes mutual annihilation with an atomic electron, gener-ally forming a short-lived (∼1 nsec) species known as positro-nium at or near the end of the positron range25—the positron-electron pair then annihilates with the emission of two 511-keV photons that travel in opposite directions (180

◦in the rest

frame). This process is much more common than the emission

of three or more photons. The finite positron range (a fewmm) and the noncolinearity of the annihilation photons (dueto nonzero kinetic energy at the time of annihilation) imposetwo fundamental limitations on the spatial resolution. The for-mer introduces a position distribution with long tails and fullwidth at 20% of maximum amplitude between 0.5–10 mm,26

which strongly depends on the positron’s energy spectrum(the radioisotope used) and the type of tissue in whichthe annihilation occurs.26, 27 The latter, annihilation photonnoncolinearity, blurs the back-to-back emission with anangular dispersion of about 0.47◦.25

The two 511-keV annihilation photons must interact in thedetectors (i.e., crystals) of the PET scanner to produce a mea-sured coincident event. In most current scanners, each scin-tillation crystal face is a rectangular shape with sides typi-cally between 4 and 8 mm, which together with the detectorring radius (r ≈ 35 cm) have a strong effect on PET spa-tial resolution.26 The multiple detectors are aligned in ringsorthogonal to the longitudinal table axis in an efficient dualcoincident photon detection geometry that for a single ta-ble position provides a typical scan length of ∼15–18 cm inthe longitudinal direction. Current clinical PET scanners havesystem spatial resolutions typically from 4 to 6 mm full



width at half maximum (FWHM) and because of partial-volume averaging have difficulty accurately quantifyingactivity in tumors with diameters smaller than about 25 mm.Tumors with diameters comparable to or even smaller thanthe system spatial resolution may nonetheless be visualizedif the tumor-to-normal tissue contrast is sufficiently large.For details on the types of current scintillation detectors,associated readout electronics, and PET scanner geometryand performance, the reader is referred to several literaturereports.28–30

Detected annihilation photons undergo energy discrimina-tion with an energy window centered around 511 keV (forexample, from 350 keV to 650 keV). A coincident event isproduced if two energy-discriminated photons are detectedwithin a timing coincidence window, usually between 6 nsand 12 ns in duration. If the two photons originate from onepositron annihilation then this event is a true coincidence.Accidental detection within the timing window of photonsfrom different annihilation events produces “random (false)events.” If one or both of the photons from an annihilationpair undergoes Compton scatter, but retains sufficient energyto lie within the detection energy window, an erroneously po-sitioned “scatter event” is produced. The rate of random coin-cidences increases with the total activity in the patient withinthe longitudinal field-of-view (FOV) of the scanner and withthe duration of the coincident timing window. The rate of scat-ter coincidences depends on the location within the body andthe patient diameter and can exceed 60% of the total detectedevents for large diameter patients.

Some older PET scanners operate in two acquisitionmodes, two-dimensional (2D) and three-dimensional (3D). Inthe 2D mode collimating septa extend between adjacent de-tector rings towards the patient and limit the detection of co-incidences only to detectors from the same ring or within asmall number of neighboring rings (typically 1 to 5). In 3Dmode, the septa are removed and coincidences are allowedbetween detectors from a large number of rings, which leadsto higher detection efficiency. Due to recent improvementsof the event processing electronics and of the reconstructionand scatter-correction algorithms, all new PET scanners oper-ate in 3D mode. Standardized procedures for determining theoptimal activity concentration, the sensitivity, the resolution,

the scatter fraction, and other parameters of the PET scannershave been developed.31, 32

The quality of PET images is affected by fundamen-tal physical and technical limitations, as described in partabove and elsewhere,29, 33 the related loss of contrast forsmall objects (partial volume effect),34–36 degradations due toattenuation,37–39 random events,40, 41 scatter,42–46 sampling,47

arc effects (nonequidistant radial sampling which leadsto nonuniformities),48 the depth of interaction,49 variable-background activity,50 and organ and patient motion.51, 52 Cor-rection techniques have been developed and progress hasbeen made in decreasing the effects of these factors on PETimage quality and accuracy.35–37, 53–59 Some of these limita-tions and their impact on the accuracy of PET segmentationof tumors are discussed in greater detail in an open accessbook chapter.60 Rapidly advancing technology and computingpower continue to improve PET scanner performance. Exam-ples of technology improvements include the introduction ofsmaller and faster responding detectors, more specialized andefficient signal processing electronics, and inclusion of res-olution and time of-flight information in the reconstructionprocess.49, 59, 61–63

Molecular imaging with the use of targeted radiolabeledtracers typically involves the following three main steps: (i)design of molecules that can target specific soluble or im-mobile cell receptors, called ligands; (ii) attachment of ra-dioisotopes that emit photons without changing the biologicalbinding properties of the target molecules; and (iii) injectionand in vivo imaging, by using high resolution and high sen-sitivity devices. PET allows quantitatively accurate detectionof molecularly targeted radioisotopes to determine tissue bio-logical properties or processes, such as commonly performedimaging of glycolysis and cell proliferation. However, imag-ing of hypoxia, gene expression for both intracellular and cellsurface receptors and proteins, tumor angiogenesis, and apop-tosis are rapidly advancing.9, 10, 64–67 With direct radioisotopeimaging, the radiolabeled material may be attached to a spe-cific molecule that selectively determines the binding locationof the labeled material. Examples include imaging of cell-surface epitopes using radiolabeled antibodies and of specificreceptors using molecularly targeted peptides (Fig. 1).65, 68

In this example, the image intensity distribution directly

FIG. 1. microPET images of three BALB/c mice implanted with EMT-6 cells and imaged with 18F-FLT (a), and 76Br-labeled antibodies with a specific bindingactivity to Sigma-II in nonblocked (b) and blocked (c) animals. Animals were imaged in supine position. Reprinted with permission from D. J. Rowland, Z.Tu, J. Xu, D. Ponde, R. H. Mach, and M. J. Welch, “Synthesis and in vivo evaluation of 2 high-affinity 76Br-labeled σ2-receptor ligands,” J. Nucl. Med. 47,1041–1048 (2006). Copyright c© 2006 by Society of Nuclear Medicine.



corresponds to the process being tracked (i.e., the numberof bound receptor sites). Alternatively, indirect radioisotopeimaging is commonly used to monitor gene expression byimaging interaction between the “reporter gene” and a com-plimentary “reporter probe.” Surrogate imaging can also beperformed, whereby processes or species can be imaged suchthat the distribution or number of a particular molecularspecies can be inferred.65

PET imaging can yield high contrast-to-noise levelsusing only a trace amount of radioactive material, on theorder of 10−7 moles or less.39 High contrast-to-noise levelsmay be limited due to nontarget (i.e., background tissues)localization of the administered radiotracer. Therefore,radioisotope molecular imaging techniques have a wide rangeof applications and are limited in their usefulness only by(1) physical properties and equipment designs that result ina finite spatial resolution and minimal detectable activity and(2) the availability and development of ligands that can carrythe radioisotope to a specific, targeted molecule and with highenough sensitivity (i.e., capability to discover the targetedfeature) and specificity (i.e., low false-positive rate) requiredto image the molecular species or process of relevance.69, 70

However, it is important to recognize that, as for systemicallyadministered materials generally, even highly targeted PETradiotracers are distributed in smaller or larger amounts to alltissues of the body.

Conventional PET relies primarily on the use of fluo-rodeoxyglucose (FDG) to image regions that are metaboli-cally active, and thus, have increased glycolysis. Uptake ofFDG may be due to abnormally high glucose metabolism(presence of tumor, inflammation, or other abnormal process),or normally high glucose metabolism (e.g., as in the brainor heart). FDG PET imaging will also exhibit high “back-ground” activities in excretory organs such as the kidneys andurinary bladder independent of their respective rates of gly-colysis. Though nonspecific for tumor and because of gen-erally increased tumor metabolism, FDG PET is particularlyuseful in staging of previously diagnosed lung, breast, colon,cervical, and head/neck cancers as well as of melanoma andlymphoma70 and for imaging nodal and distant metastases.Both metabolic level (i.e., signal intensity) and changes inglucose metabolism can be used for tumor diagnosis and eval-uation of treatment response. The standardized uptake value(SUV) is an imaging parameter that is sometimes used as athreshold above background for determining a positive diag-nosis for metabolic activity, and thus, malignancy.71 There isnot a universal definition for SUV although it is generallydefined as

SUV = activity(MBq)/mltissue

administered activity(MBq)/gbody mass. (1)

The SUV is calculated either on a voxel-by-voxel basis orover a region-of-interest (ROI). Also, in alternative forms theSUV calculation may use the lean body mass or body sur-face area. Typically, either the maximum or mean value of theSUV in a ROI is used clinically and an SUV of 2.5 or greateris commonly used as the diagnostic threshold indicating ma-lignancy for FDG PET.72, 73 While tumor SUV generally cor-

relates with the degree of malignancy, such use of SUV ishardly unambiguous—observed SUVs can be affected by im-age noise, coarse image spatial resolution, other concurrentnonspecific processes (e.g., perfusion), and user biased ROIselection. An AAPM Task Group is currently working to de-fine the use of SUV for PET-based segmentation. In summary,SUV is a clinically useful quantity for the parameterization ofPET images, but care has to be taken with the interpretationof the results.74

Other applications of PET for molecular imaging includecerebral blood flow (perfusion) using 15O H2O, tumor hy-poxia (Fig. 2) with 18F fluoromisonidazole75 and 64Cu-ATSM(diacetyl-bis(N4-methylthiosemicarbazone)),76 and cell pro-liferation with 11C thymidine, 18F fluorothymidine (FLT), andother radiolabeled thymidine analogs.77 Although all aminoacids have been radiolabeled, methionine and tyrosine havebeen focused upon with PET imaging.69 Furthermore, syn-thetic amino acids such as 18F-fluoro-aminocyclobutyric acid(FACBC) have also been used for tumor imaging.78 Other ad-vanced PET applications in oncology include imaging cellpermeability, deoxyribonucleic acid (DNA) synthesis, tumorreceptors, and monitoring of the distribution of and responseto therapeutic agents.70

The hybrid PET-CT scanner79 rapidly replaced PET-onlyscanners and is contributing to the radiation oncology com-munity because it provides both a CT image with the neededelectron density information for accurate radiation dose cal-culations and an accurately coregistered biomolecular PETimage of the patient in the treatment position (Fig. 3). Soft-ware tools available in most commercial radiation treatmentplanning systems allow display and manipulation of the fusedanatomic (CT) and biomolecular (PET) images.

3.B. Single photon emission tomography

Single photon emission computed tomography (SPECT)is a nuclear medicine tomographic imaging techniquethat employs the emission of x-rays and noncoincidentgamma-rays by radiolabeled agents. Those radiolabeledtracers are injected intravenously in the human body andconcentrate in an organ or structure of interest. Radiolabeledtracers are produced by the combination of gamma-emittingradioisotopes and a ligand that is designed to bind to certainmolecules thereby targeting a particular anatomic area ormetabolic or other biologic process. The emitted photonsmay pass through a collimator and the detector producesa 2D image of the tissue localization of the radiotracer.Compared to PET, using a collimator has the significantdisadvantage of reduction in sensitivity. SPECT is a tomo-graphic technique and produces 2D cross-sectional imagesat multiple angles around the subject that can be furthermanipulated for multiplanar or 3D volumetric reformats asneeded. In contrast to PET, which detects time-coincidentphoton emission of positron-electron annihilation occurringin very close proximity to positron generation, SPECTprovides direct detection and measurement of overalltissue photon emission that inherently possesses less radi-ation event localization information and therefore results



FIG. 2. Two separate examples of hypoxia PET imaging (all images are transverse sections). (a)–(c) Brain tumor imaging using (a) T1-weighted and (b) T2-weighted MR imaging followed by (c) 18F-fluoromisonidazole PET hypoxia imaging. (Arrow indicates a hyper-intense region surrounding the hypo-intensecenter of the tumor) (Images courtesy of JD Bourland, Wake Forest School of Medicine). (d) and (e) Imaging of head/neck cancer using (d) CT and (e) 60Cu-ATSM PET hypoxia imaging. From K. S. Chao et al., “A 1160 novel approach to overcome hypoxic tumor resistance: Cu-ATSM-guided intensitymodulatedradiation therapy,” Int. J. Radiat. Oncol., Biol., Phys. 49, 1171–1182 (2001). Copyright c© 2001 by Elsevier Inc.

FIG. 3. Fused fluorodeoxyglucose (FDG) PET/CT images in the (a) transax-ial, (b) sagittal, and (c) coronal planes of a patient with lung cancer. CT wasacquired at 45 mAs, 120 kV, pitch ∼1.The PET scan was acquired 60 minafter the injection of FDG (3D acquisition at 3 min per bed position). A largetumor in the left lung is apparent (arrow). Also present is malignant lym-phadenopathy in the mediastinum (arrow head) and metastatic lesions in theliver (dashed arrow). Image provided by Dr. Bradley Kemp, Department ofMedical Physics, Mayo Clinic, Rochester, MN.

in poorer image spatial resolution, typically several millime-ters to 1 cm.

The main detector component of a clinical SPECT systemis the Anger gamma camera.80 For each incident photon, agamma camera collects signals from a number of photomul-tipliers (usually 50–70) that cover the entire field of view.Those signals are multiplexed by using a resistive networkand only four position signals are derived, which are usedto calculate the X and Y coordinates of each photon. Moredetails on calculating the coordinates of the incident pho-tons are available in the IAEA Human Health Series ReportNo. 6.81

When the gamma camera is rotated around the object to beimaged, projection data are obtained and tomographic SPECTimages are mathematically reconstructed.82 Typically, pro-jections are acquired every 3◦–6◦ while the camera rotatesaround the object of interest. A full 360o rotation is nec-essary to produce a reconstructed image of optimal quality.Depending on the imaging application 60 to 120 projectionsmay be acquired, for example, in cardiac imaging, wherethe camera rotates 180o from approximately the right ante-rior oblique to the left posterior oblique to minimize the pro-nounced attenuation and scattering effects associated with ra-diation emitted latero-posteriorly from the heart. Correctiontechniques and reconstruction algorithms can compensate forthose effects, but their clinical value and acceptance are stillbeing addressed.83, 84 Acquisition of each projection typicallylasts 15–20 s, which yields a total scanning time of around



20 min. Scanning times may take as long as 40 min for lowadministered activities and/or high-energy photon emitterssuch as gallium-67 and iodine-131. Of interest, acceler-ated acquisitions can be accomplished by using multiheadedgamma cameras.

In the Anger camera a lead or tungsten collimator, usu-ally with hexagonal parallel holes, determines the photonsthat will be accepted and measured. The collimator is themain limiting factor of the resolution and sensitivity of clin-ical SPECT systems and its optimization remains an openissue, whereas in dedicated small animal imaging systemsthe multipinhole focused collimators appear to be the designof choice.85 A scintillator converts each photon, after pass-ing through the collimator with initial energies that vary be-tween 71 keV (Thallium-201) and 364 keV (Iodine-131), tothousands of optical photons with energies of several eV.86, 87

This conversion-amplification task is performed by a combi-nation of a scintillator with a photomultiplier tube (PMT). Thescintillator is coupled to the PMT, usually through an opticallight-guide such as silicon fluid, grease, or Lucite light pipes.Optical photons strike the PMT’s photocathode, which emitsphotoelectrons. By applying a high, usually negative voltage(∼1000 V) to a multidynode stage, the photoelectrons aremultiplied, forming an electron cloud that hits the PMT’s an-ode. This charge is detected and further processed by acqui-sition electronics. Although all clinical SPECT cameras usescintillators made of sodium iodide doped with thallium (NaI-Tl), several new materials including cesium iodide doped withthallium (CsI-Tl), cesium iodide doped with sodium (CsI-Na), and lanthanum bromide doped with cerium (LaBr3:Ce)may offer improved light output, energy resolution or PMTcoupling.87 Solid state detectors, which allow direct photondetection and have greatly improved energy resolution, are arelatively recent technological alternative to the conventionalscintillator-PMT module.88 Energy discrimination by solidstate detection can be crucial in novel applications where mul-tiple radiotracers are administered and detection of gammaray emission of different energies is necessary in order todifferentiate tissue colocalization of different radionuclides.89

Even more importantly, however, may be the improved scat-ter rejection, and therefore improved contrast and even spatialresolution, resulting from superior energy resolution.

There are two major advantages of SPECT compared toPET. The first is the existence of radiotracers that are based ona variety of radioisotopes with different energies and their rel-atively long half-life that makes them easily accessible to rou-tine clinical practice. The second is radioisotope availabilityfor academic research, even at centers that are far away fromcyclotron facilities. For an extensive description of the variousradiotracers available for SPECT imaging and their distinc-tive properties the reader is referred to the review paper byLoudos et al.90 Currently, the radionuclides most frequentlyused for SPECT radiolabeling include 99mTc, 111In, 67Ga, 131I,and 201Tl.91–93 Tc-99m and In-111 are gamma emitters thatare widely used for research purposes due to their availabil-ity, suitable half-lives, and relative ease of attachment to smallmolecules of interest.94, 95 Endogenous ligands such as pep-tides, antibodies, hormones, and selectins can be relatively

easily labeled for SPECT imaging.96, 97 Also, because of theirsize these molecules diffuse slowly into tissue and have slowclearance from blood, which can be of the order of hoursor even days. Thus, the long half-lives of commonly usedSPECT isotopes makes imaging possible for slow biologicalprocesses (e.g., cell division, infection, and inflammation) andfor validating the biodistribution of slowly localizing thera-peutic radiopharmaceuticals.

Myocardial, brain, thyroid, and whole-body bone SPECThave become routine in the clinic. Myocardial SPECT in-cludes cardiac-gated acquisition to provide quantitative pa-rameters of myocardial function and perfusion such as heartmuscle thickness, myocardial contractility, left ventricularejection fraction, and cardiac output. Other specific appli-cations are being developed.98 Dynamic renal scintigraphyshows that SPECT has the potential to provide valuable dy-namic information. These studies are limited to the preclini-cal level and their clinical merit is to be determined.99 SPECTimaging of inflammatory properties of vulnerable atheroscle-rotic plaque is an active area of research, in order to pre-emptively identify high-risk coronary cases and individual-ize patient treatment.100 In the near future, innovative ap-plications of SPECT imaging may include radioactive label-ing and noninvasive in vivo tracking of stem cell traffickingand tissue homing in the interest of targeted cardiovascularstem cell therapies for myocardial infarction and therapeuticangiogenesis.101, 102

SPECT has an increasing role in oncology due to a numberof novel radiopharmaceuticals that are designed for specifictumors. For example, 99mTc-HYNIC Octretoide (99mTc-OCT)images somatostatin receptor expression and distribution,while 99mTc-Glucarate (99mTc-GLU) has increased tumoralaffinity.103–105 SPECT tracers are being used for functionalimaging and anatomical characterization of brain tumors,106

and in a preliminary evaluation have been used for tumor di-agnosis, staging, monitoring of therapy outcomes, and detec-tion of recurrence in lung cancer and malignant lymphoma.107

In early work, SPECT was used to evaluate pre- and post-treatment lung function and thus plays a role in the optimiza-tion of lung radiation treatment plans.108 In a similar manner,reproducible molecular imaging of tumor metabolic activitywill allow better localization of tumor regions and early as-say for the radiation response of tumors and normal-tissuesthereby increasing overall treatment safety and efficacy(Figs. 4 and 5).109

In basic oncology research, numerous preclinical studieshave developed protocols of molecular imaging of tumor cellreceptors, cell apoptosis, hypoxia, and angiogenesis.110–113

SPECT has been used to image integrin ανβ3 expression,VEGF, extracellular matrix, activated endothelial cells, andmatrix metalloproteinase, all of the latter being implicatedin tumor neovascularization.114 SPECT also has the uniqueability to probe two or more molecular pathways simultane-ously by dual-isotope/dual-energy imaging;115 thus, differentorgans or functions can be monitored simultaneously.116 Re-sults from simultaneous imaging of Tc-99m and I-123, andof Tc-99m and Ga-67 are available.117 Multimodality SPECTimaging, usually combined with CT (hybrid SPECT-CT),



FIG. 4. 131I SPECT study in a female patient with follicular thyroid cancer, presenting with relapse of the disease 2 years after thyroidectomy and radioiodineablation of the surgical remnant. Images were acquired 6 days after a new therapeutic administration of 4.5 GBq (150 mCi) of 131I. Acquisition settings were:energy photopeak 364 keV; energy window 20%; angular range 3600; 120 projections; and 20 s per projection. Tomographic reconstruction was accomplishedby an iterative algorithm (OSEM, 2 iterations, 10 subsets). Consecutive coronal slices are presented, which show local cancer recurrence (solid arrows), cervicallymph node metastases (dashed arrows), and multiple lung metastases (arrowheads). The high contrast of the images is due to the high administered activity andthe high tumor-to-background 131I uptake ratio, particularly in the neck.

has already evolved to standard practice in several fields ofoncology, cardiology, and neuropsychiatry in order to com-bine functional and anatomical image acquisitions.118 HybridSPECT-CT imaging can achieve accurate coregistration ofimage data and improve diagnostic evaluation of indetermi-nate lesions detected on routine cross-sectional imaging ofoncologic cases.119 SPECT is also expected to become a keyelement of future molecular imaging approaches, along withgenomics, proteomics, metabolomics, and radiogenomics,contributing to the convergence of tissue-specific diagnosticsand tissue targeted therapeutics, i.e., theragnostics.120

3.C. Magnetic resonance imaging

The origin of the signal in magnetic resonance is the mag-netic dipole moment, μ, that results from those nuclei inwhich the net angular momentum is nonzero. In the presenceof an external magnetic field (B0) μ will undergo precessionabout B0 with an angular frequency, ωo, defined by the Lar-mor equation, ωo = γ B,0 where γ is the nuclide-specific con-stant gyromagnetic ratio. At 1.5 T, this frequency is approx-imately 64 MHz for protons, placing it within the radiofre-quency (RF) portion of the electromagnetic spectrum.

FIG. 5. 111In-Octreotide (OcteoScanTM) SPECT study in a patient with suspected neuroendocrine tumor of the pancreas. Images were acquired 24 h afterradiotracer administration of an activity of 200 MBq. Acquisition settings were: energy photopeaks 171 and 245 KeV; window 20% around each photopeak; an-gular range 3600; 120 projections; and 30 s per projection. An iterative algorithm (OSEM, 2 iterations, 10 subsets) was employed in tomographic reconstruction.Selected transverse, coronal, and sagittal slices are presented, which show abnormal focal tracer uptake at the site of the pancreatic tumor (red cross). Normaltracer accumulation in the spleen (solid arrow), the lower part of the liver (dashed arrows) and the urinary bladder (arrowhead) is also demonstrated.



Signal formation, signal-to-noise ratio (SNR) and contrastare a result of the complex interaction of μ with its local en-vironment. MR image formation is achieved by the applica-tion of RF pulses followed by linearly varying magnetic fields(gradients) that spatially encode the MR signal.121 Image con-trast is determined by the relative distribution of tissues ofdiffering T1 and T2 relaxation times and the choice of MRimaging parameters. The T1 time constant describes the rateof energy loss to the lattice structure of the material (i.e., spin-lattice) while the T2 constant describes the loss of phase co-herence due to spin exchange between nuclei (i.e., spin-spin).A third time constant, known as T2*, includes the loss of sig-nal due to macroscopic magnetic field inhomogeneities and isrelated to the T2 constant by the relationship 1/T2* = 1/T2 +γ�B0 where �B0 describes the inhomogeneity. Parametersthat emphasize T1 or T2 differences generate the so-called T1-and T2-weighted images, respectively. A third type of contrastin which neither T1 or T2 relaxivity is emphasized but ratherthe relative difference in tissue proton density can also be ac-quired. These images are known as proton density-weightedimages.

The SNR in an MR image is in large part determined by thefield strength of the main magnetic field, B0, which in com-mercially available devices is 1.5 T (most common) or 3.0 T(gaining wider acceptance, particularly for neurologic imag-ing applications). It is noted that within the United States,the US Food and Drug Administration (FDA) has granted510(k) clearance for commercial distribution of whole bodyMR scanners at field strengths of 3.0 T or less. Human MRimaging with field strengths above 3.0 T is restricted to re-search institutions as defined by approved Institutional Re-view Board (IRB) protocols, with maximum field strengths of8.0 T for adults, children, and neonates older than one monthand 4.0 T for neonates less than one month. Research imag-ing with humans at field strengths beyond these limits requirespecific permission by the FDA in the form of an investigativedevice exception (IDE).

MR spectroscopic imaging (MRSI), a variation of MRspectroscopy (MRS), is an in situ method of determiningthe relative concentrations (i.e., MRS) and spatial distribu-tion of metabolites containing the nuclide of interest. 1H isthe most common nuclide imaged using MRSI methods dueto its relative abundance compared to other nuclei and hasbeen demonstrated to detect, quantify, and differentiate neo-plastic disease processes in the brain, breast, and prostate.122

Brain tumors are characterized by a loss of N-acetylaspartate(NAA), elevated levels of choline-containing compounds andlactate (in certain tumors), and reduced levels of creatine.Each of these chemicals provides an important signature oftumor growth and behavior: decreased NAA levels are as-sociated with decreased neuronal density in the tumor, ele-vated choline is indicative of altered-membrane phospholipidmetabolism, increased lactate levels result from increasedanaerobic metabolism, and decreased creatine levels are in-dicative of altered bioenergetic pathways.122 MRS/MRSI ofbrain cancers has been used for the purposes of diagno-sis/tumor grading,123 radiation treatment planning124 and re-sponse assessment, following administration of therapy125 as

well as information on prognosis and survival.122 For MRSIof the prostate, the key biochemical changes include a three-fold reduction in citrate concentration and a two-fold in-crease in choline concentration compared to healthy periph-eral zone regions.126 The peak area ratios of choline/citrateand (choline + creatine)/citrate appear to be reliable indica-tors of cancer in the prostate.126 Strong linear correlations be-tween Gleason grade and both a decrease in citrate and in-crease in choline concentrations have been shown,122 sug-gesting that MRSI can be used as a metric for the assess-ment of radiation therapy efficacy. In breast cancer, increasesin the levels of choline and its metabolites associated withcell membrane synthesis and turnover resulting from malig-nant cell growth can be monitored using MRS127 to determinethe presence/absence of cancer. 1H MRS has also been inves-tigated for diagnosis of neoplasms in a variety of other or-gans including liver,128–131 bone,132 and salivary glands133 andfor assessing treatment response in head and neck cancers,134

thyroid,135 and melanoma.136 MRSI of nuclei other than hy-drogen, such as 13C, 19F, 23Na, 31P, 35Cl, and 39K, is also pos-sible. 31P is particularly significant because its presence canbe used to elucidate intracellular pH and ATP homeostasis.137

However, the relative abundance of these nuclei is of the orderof 10−4 to 10−7-fold less than that of 1H.

Acquisition of MRSI data can be performed using chem-ical shift imaging techniques.138 Figure 6 shows two spectrafrom voxels within the prostate overlaid onto a T2-weightedimage of the gland. The first spectrum, which is located onthe left side of the gland within the peripheral zone, identifiesa high citrate peak with relatively low choline and creatinepeaks typical of healthy prostatic tissue. The second spec-trum, taken from a different slice including both the centraland peripheral zones, shows an absence of citrate with an ele-vated choline peak, suggestive of malignancy. Notice also thatin this slice the spectroscopic voxel is located over a region ofdecreased signal on the T2-weighted images and shows thecharacteristic signal decrease in those regions where cancer ispresent.

MRSI is technically challenging in comparison to con-ventional MR imaging methods and as such its integrationinto routine clinical diagnostic and radiation therapy plan-ning has been slow. MRSI data do not directly lend them-selves to familiar grayscale images, which are typically rep-resented in pseudocolor-scale, and no commercial radiationtreatment planning systems exist at this time that can readilyutilize MRSI data. MRSI exhibits rather coarse spatial resolu-tion, with typical MRSI voxel sizes on the order of 0.1–1 cc.Smaller sampling sizes are possible, but require longer imag-ing times and/or higher magnetic field strengths.139

Perfusion is the mechanism by which nutrients and oxy-gen are delivered to and waste products are removed fromcells.140 This process can be visualized by MR using a perfu-sion contrast agent that can be either endogenous or exoge-nous. Blood is an effective endogenous contrast agent andperfusion imaging with blood as the contrast agent is per-formed by acquiring two data sets: the first in which the mag-netization of the blood on the arterial side of the imagingvolume is nulled or saturated (measured signal is set as the



FIG. 6. MR 1H spectroscopic profiles characteristic of (a) normal and (b) cancerous prostate tissue. T2-weighted axial images are used to identify the region ofinterest from which the spectra is sampled (square region) with the associated metabolic profile shown adjacent. Normal peripheral zone tissue is characterizedby a high citrate and low choline concentrations as seen in (a). In contrast, the large hypo-intense mass (arrows) seen on the T2-weighted image in (b) shows amarked elevation of choline and no identifiable citrate peak. Metabolite locations are identified by their relative frequency shift from the water resonant frequencyin units of parts per million. Metabolite peaks are measured in relative units. Images courtesy of Dr. Akira Kawashima, Department of Radiology, Mayo Clinic.

reference) and a second in which fully magnetized arterialblood is allowed to flow into the same volume (control). Thedifference between the control and reference data provides anestimate of the blood delivered to the imaging volume; thistechnique is known as arterial spin labeling.141 Exogenouscontrast agents are substances that change the relaxivity of theblood and are administered intravenously. The agent is typi-cally injected as a bolus and is imaged with rapid MR tech-niques (echo planar imaging (EPI)) to detect the wash-in andwash-out of the agent into and out of the tissue of interest,respectively. This approach is known as first pass perfusionimaging or dynamic contrast-enhanced MRI (DCE-MRI).142

A variety of tumor specific data can be obtained fromperfusion MR images143 (Fig. 7). Differences in such im-ages between tumors and normal-tissue are primarily dueto variations in the vasculature of these respective tissues.T1-weighted perfusion images using DCE-MRI techniqueshave been used to characterize lesions in the breast144 andmuscle.145 In breast cancer, rapid uptake of low molecularweight contrast agents has been shown to be an indicator of atumor and represents the most commonly used application ofdynamic contrast-enhanced perfusion imaging for cancer di-agnosis. However, fibroadenomas also often exhibit enhancedvisualization by DCE-MRI, reducing its specificity for ma-lignant disease in the breast.143 Early enhancement in DCET1-weighted MR perfusion images has also been used for thestaging of gynecological, bladder, and prostate cancers as wellas for monitoring treatment efficacy following administrationof chemotherapy.143 Qualitative changes in T2*-weighted per-fusion MR images have also been used to differentiate livermetastases, hemangiomas, and hepatomas146 as well as pro-vide a method for differentiating carcinomas and fibroadeno-mas in breast tissue.147 Though MR perfusion imaging hasbecome more common for diagnostic purposes, this imagingapproach is an area of research for the radiation oncologycommunity.

In diffusion MRI, the relative mobility of water withinthe imaging volume is the basic property from which signalcontrast is derived. Any unbound molecule (water or other)will undergo random translational, or Brownian, motion. Un-der the influence of a magnetic field gradient, the MR sig-nal, S, from a molecule undergoing diffusion will decreaseaccording to the monoexponential relationship, S = Soe−bD

where So is the signal in the absence of diffusion, “b” isa factor proportional to the integral of the applied diffusion

FIG. 7. Relative cerebral blood volume (rCBV) measurements obtainedfrom an MRI perfusion imaging sequence. Asymmetric perfusion distributionbetween opposing hemispheres indicates an unequal distribution of blood.The relative increased blood flow around the tumor periphery (arrow) indi-cates breakdown of the blood-brain barrier while the lack of signal within thecentral zone of the tumor indicates necrosis/calcification. rCBV values colorcoded in terms of the number of standard deviations from the mean cere-bral blood volume within a normal (i.e., reference) region of the brain. Imageprovided courtesy of Dr. Kirk Welker MD, Department of Radiology, MayoClinic.



sensitizing gradient waveform(s) over their duration (s/mm2)and D is the diffusion coefficient (mm2/s). Acquiring an MRIimage with and without the application of this diffusion sen-sitizing gradient provides a measure of S and So, allowing anestimate of D.

The extent of diffusion of water in vivo is highly depen-dent upon a variety of factors that include fluid viscosity, cel-lular permeability, active transport mechanisms, and the mi-crostructure of the local environment.148 As a result of thesensitivity of diffusion MR imaging to these tissue proper-ties, diffusion-weighted MR imaging has been used for a va-riety of oncologic applications. MRI-derived cellularity dif-ferences among tissues have been correlated with diffusion-weighted MRI and have been associated with the onset ofneoplastic changes,149 the ability to differentiate tumor frombenign or necrotic tissue,148 and tumor grade.150 The seriallymonitored apparent diffusion coefficient (the measured diffu-sion value) has also been shown to correlate with response tochemotherapy151 as well as to be a potentially useful tool formonitoring therapeutic response to radiation therapy of braincancer.152 Diffusion tensor MR imaging, which quantifies dif-fusion anisotropy as a second-rank tensor within each voxel,can be used to track the maximum diffusion value alongnerve fibers, providing so called tractography images, and hasbeen used extensively for mapping white matter tracts in thebrain153 (Fig. 8). Fiber tracking using diffusion tensor infor-mation is also providing important treatment planning infor-mation for both surgical and radiation therapy approaches, al-lowing visualization of the spatial relationship between thesetracts and tumor and their distortion by the tumor.154–158 Be-

cause of the technical challenges and dependence upon userinput, tractography images are not currently considered suffi-ciently reliable to be useful for routine radiation therapy plan-ning applications.

Functional MR imaging (fMRI) provides information onthe hemodynamic and metabolic changes in the brain fol-lowing neuronal excitation.159 It has been shown that dur-ing excitation cerebral blood flow can increase between ap-proximately 30%–50%. The increased perfusion provides aconcomitant increase in oxyhemoglobin, increasing the rel-ative concentration of oxygen in activated regions. How-ever, oxygen consumption in these regions increases by ap-proximately 5% as the baseline state of aerobic metabolismswitches to the active state of anaerobic glycolysis.160 Thedisassociation of oxygen supply and consumption forms thebasis of the contrast mechanism used in fMRI and is knownas blood oxygen level-dependent (BOLD) contrast.161 Themagnitude of the BOLD response is determined by the rela-tive concentrations of deoxyhemoglobin (paramagnetic) andoxyhemoglobin (diamagnetic) within the blood.162 Param-agnetic substances produce local magnetic field inhomo-geneities which result in susceptibility induced dephasing andsignal loss.121 T2*-weighted gradient echo imaging sequencesare used to quantify the signal difference by acquiring twodata sets, the first in the resting state in which the brain issaturated with deoxyhemoglobin and the second followingneurological activation when the inflow of oxyhemoglobin isgreatest. Because it is noninvasive in nature and avoids the useof ionizing radiation, fMRI has been used for the assessmentof neuronal function pre- and post-treatment and to identify

FIG. 8. White matter fiber tracks derived from diffusion tensor imaging superimposed onto an axial T2-weighted image demonstrating major motor cortexfibers traveling inferiorly towards the brain stem. The corona radiata/centrum semiovale are the fan shaped superior fibers seen radiating superiorly from theselected T2-weighted axial magnitude image (a)–(c). These fiber tracks coalesce inferiorly into fiber tracts that include the corticospinal tracts. Comparison ofthe tracks within each hemisphere allows visualization of the impact of the tumor on these critical neurological structures. (b) View of left hemisphere fibertracks illustrating fiber tract displacement (solid white arrow) due to the presence of the tumor. (c) View of the right hemisphere tracts showing their normal,unperturbed distribution (dashed white arrows). Fiber tracts are generated by identifying seed voxels and tracing of their connectivity based on the direction ofhighest diffusivity based on diffusion tensor information. Image provided courtesy of Dr. Kirk Welker MD, Department of Radiology, Mayo Clinic.



FIG. 9. Hybrid image showing areas of functional activation (colored pixels) superimposed onto an echo-planar imaging (EPI)-based spin echo magnitudeimage. Areas of functional activation are generated by measuring signal changes arising from blood oxygen level dependent (BOLD) contrast between noactivation (i.e., resting state) and task-induced activation. Broca’s area (arrow in image 14) and Wernicke’s area (arrow in image 18)—anatomical regions of thebrain associated with language—show hemodynamic activation allowing spatial identification of these locations relative to the tumor. Image provided courtesyof Dr. Kirk Welker MD, Department of Radiology, Mayo Clinic.

brain regions to be avoided (Fig. 9) with either surgical or ra-diosurgical treatment.163, 164

For exogenous contrast, and because of its paramag-netic properties, solutions of chelated organic Gadolinium(Gd) complexes are used as intravenously administered MRIcontrast agents.165, 166 To avoid Gd3+ toxicity, Gd contrastagents are encased in a chelate molecule. Gd-based con-trast agents reduce tissue T1 and T2 relaxation times andare used to increase tissue contrast resolution. A new classof Gd contrast agents is currently under investigation asspecific agents for molecular imaging. These agents differfrom Gd-DTPA complexes by the inclusion of functional el-ements into the molecular complex that can recognize andbind to specific molecules or in relation to specific molecu-lar processes occurring on or within the cellular membrane.Like targeted drug therapy, these molecular constructs arecharacterized by a functional unit that provides the speci-ficity necessary to identify the molecular process, a molec-ular “vehicle,” and the Gd ion(s). Ideally, the concentrationof Gd at each cell should be sufficient to identify them in-dividually. However, such a scenario requires voxel dimen-sions to be of the same resolution as the cell’s diameter,which is not achievable on clinical, whole body MR imag-ing systems. The functional unit can serve to attach themolecular structure to a specific cell receptor or provide amechanism by which Gd can be delivered directly acrossthe cellular membrane to accumulate within the cytoplasmof the cell. Aime et al.167 and Sosnovik and Weissleder168

have presented reviews of these processes. Such ap-proaches fall under the current scope of nanotechnology MRimaging.169–171

3.D. Optical imaging

Optical molecular imaging is based on detecting visibleand infrared photons after they are transmitted through bio-logical tissues. The resulting images are a result of photonemission, reflectance, absorption, scattering, and phase-shift.The advantages of optical imaging are its noninvasiveness,real time observation and acquisition, ease of use, low cost,and minimal toxicity to biological systems. Since energy lev-els of atoms emitting infrared and optical photons are low,from ∼1 to ∼10 eV, path-lengths (i.e., penetration depths) intissue are relatively short (several cm at most). Thus, opti-cal imaging tends to be used for relatively thin geometries,and has been more widely used for in vitro measurementsand surface or near-surface in vivo imaging of small animals.Major optical imaging techniques are bioluminescence, flu-orescence, and diffuse/coherent tomography,67 ordered fromhighest to lowest sensitivity.172

Bioluminescence imaging (BLI) uses biologically pro-duced photon emissions. The predominant technique usesfirefly luciferase, the enzyme that is responsible for pro-duction of the yellow-green light emitted by fireflies173, 174

(Fig. 10). The luciferase gene is cloned and spliced intotarget DNA of specific cells ex vivo and the cells then in-fused into the study subject. Other luciferases, such as re-nilla luciferase, are also used.175 With a biochemical energysource and oxygen, luciferase catalyzes reactions that resultin yellow to green photon emission (approximately 500 nm).Self-contained BLI commercial systems are available forsmall sample sizes. Uses of BLI include imaging ofin vivo gene expression, monitoring of protein–protein



FIG. 10. Approximate correlation between BLI output and tumor size. (a)Ef-luc BLI of PDGF-induced oligodendrogliomas. Luciferase activity intumor-bearing Ef-luc N–tv-a transgenic mice. (b) Whole mount histologicanalysis of the brains from the same mice as were imaged in (a). Tumorsize roughly correlates with the amount of emitted light. Reprinted by per-mission from L. Uhrbom, E. Nerio, and E. C. Holland, “Dissecting tumormaintenance requirements using bioluminescence imaging of cell prolifera-tion in a mouse glioma model,” Nat. Med. 10, 1257–1260 (2004) (Ref. 237).Copyright c© 2004 by Nature Publishing Group.

interactions, tracking cancer cells, and whole body imagingof small animals.173, 174, 176 Also, in vivo growth and responseof tumors and metastases before and after chemotherapy havebeen studied for prostate and other cancer models.177 BLI re-mains largely a planar (or 2D) imaging modality and recon-struction of the 3D in vivo distribution of bioluminescenceis challenging because of the numerous and complex opticalphoton interactions in tissue. Bioluminescence tomography(BLT) is being pursued, however, offering the possibility ofproviding reliable depth information and greater quantitativeaccuracy.178–180

Fluorescence techniques can be divided into optical andnear-infrared (NIR) fluorescence imaging;172 green fluores-cent protein (GFP) is the most widely used of the opticalfluorophores. These techniques can be implemented usinggeneral-purpose charge-coupled detector (CCD)-based cam-eras and data acquisition and processing techniques that al-low deconvolution of the abundant tissue autofluorescencefrom that of the signal fluorescence of the specific fluorophorebeing imaged. GFP imaging has been developed based onthe discovery and isolation of that protein from jellyfish andother iridescent creatures.181, 182 Organic fluorophores havebeen used for NIR imaging. The wavelengths of light for GFPimaging range between 380 nm and 780 nm while those forNIR imaging range between 700 nm and 1000 nm.176 Usesof GFP include imaging of in vivo gene expression, mon-itoring of protein–protein interactions, and tracking of pro-tein populations.67, 182, 183 The GFP gene is inserted into a ge-netic construct which in turn is incorporated into the DNAof the cells of interest. In situ, fluorescence is stimulated byan external excitation visible-light source of the appropriate

wavelength. Absorption at 489 nm results in fluorescence at508 nm without any exogenous agent being administered—the product of the cloned GFP gene leads to fluorescence. Theresulting emission intensity pattern can be digitally capturedand displayed with a desired grayscale or color-scale, indi-cating the distribution or intensity of protein activity. GFPhas been modified to produce blue, red, yellow, and cyanversions (BFP, RFP, YFP, and CFP, respectively) throughwavelength shift mechanisms due to amino acid substitutionsresulting in mutant variants.182, 183 The use of such multiplediscrete wavelengths allows discrimination of multiple targetsor processes.

Diffuse optical tomography (DOT) is based on NIR lightthat is transmitted through tissue and detected using an arrayof sources and detectors distributed over a large region.172, 184

The NIR light is used to probe the interior of the body forphysiological changes and can provide functional data thoseare derived due to the light absorption spectra of specificmolecules. DOT can provide either scalar- or vector-basedmeasurements. Scalar methods measure the optical flux ex-iting the tissue and the photon path length. The vector meth-ods measure both the optical flux magnitude and the averagepropagation delay using either a time-domain or frequency-domain system. Time-domain systems use picosecond-wideoptical pulses and time-gated photon counting detectors;frequency-domain systems use a RF modulated light source,photomultiplier tubes or fast photodiodes, and RF phase de-tectors. The maximum imaging depth of DOT is about 1.5 to2.0 cm.185 The advantage of using frequency-domain systemsis a greater SNR compared to that of time-domain systems.DOT has been used to image the molecular function of breastcancer.184

Optical coherence tomography (OCT) uses backscatteredlight photons and a Michelson interferometry detector systemto produce 2D images of the structure of superficial layers oftissue.186–188 The principle of OCT is analogous to that of ul-trasound B-mode imaging except OCT uses light instead ofacoustic waves to differentiate tissue composition by measur-ing the light reflectance. The technique has been described as“an optical biopsy,”186 since OCT can produce near-histologicimages without excision, with a spatial resolution of 1 to15 μm. Due to photon absorption and scattering, its samplingdepth is limited to within 2–3 mm of the surface in tissues.The 2D images can be assembled to construct a 3D imageset. Originally developed for ophthalmology, OCT is beingapplied to cancer diagnosis and tissue characterization, forinstance, for optical imaging of larynx cancer, and may beable to perform imaging at the cellular level.189, 190 OCT isan example of an imaging technique that acquires signals thathave a molecular origin. Thus, it can provide novel anatomi-cal information at a scale finer than that given by CT. Poten-tial future applications of OCT in radiotherapy include refinedtarget and critical organ definitions for treatment planning forareas reachable by endoscopy, after registration to the plan-ning CT.

Recent advances in the development of nanoparticleimaging probes have introduced semiconductor-based quan-tum dots (QDs) for in vivo molecular and cellular optical



imaging.191 Quantum dots are luminescent nanocrystals com-posed of atoms from groups II–VI or III–V elements in the pe-riodic table.192 These 2- to 6-nm size QDs provide a promis-ing alternative as optical contrast agents to organic dyes andfluorescent proteins. The advantages of using QDs are tunableemission from visible to infrared wavelengths, large absorp-tion coefficients across a wide spectral range, very high levelsof brightness, photostability resistant to metabolic degrada-tion, and small size.193 Bioconjugated QDs offer new pos-sibilities and challenges for highly sensitive optical imag-ing of molecular targets in genes, proteins, living cells, andanimal models and for early cancer detection, tumor celltracking in real time, and monitoring of cancer treatment inhumans.194

3.E. Ultrasound

Developments of ultrasound for molecular imaging in-clude the characterization of tissues through spectral analysis(i.e., the spectrum of flow velocities are mapped vs time) andthe use of specialized contrast agents that enhance imagingfor a specific biological structure or process. Reflected ultra-sound carries spectral information that varies with structuresand with molecular components that comprise tissues. Tissueidentification and distinction of normal and abnormal tissueshas been and continues to be investigated.195–197 Conventionalultrasound imaging provides structural information.198, 199 Inaddition, ultrasound contrast imaging offers functional in-formation such as perfusion status, tumor viability, andmolecular expression in intravascular targets. Ultrasound con-trast imaging methods include the use of small gas-filledbubbles (microbubbles) distributed to tissues via the vas-cular system.195 Microbubbles provide contrast due to theechogenicity of the microbubble gas or its containing shell.Molecular specificity is ensured by the attachment of anti-bodies, peptides, and other ligands to the microbubble sur-face. Microbubbles with a stabilizing shell are now FDA-approved. Microbubble diameters range from 1 to 10 μm andshell thicknesses from 10 to 200 nm.200 The lifetime of mi-crobubbles depends on the characteristics of the gas inside,the bubble shell composition, and the acoustic waveform (in-tensity and frequency) striking the microbubbles. Ultrasoundcontrast can also be provided by the use of nonbubble emul-sion contrast agents.201

Major applications of ultrasound-based contrast imagingare blood vessel detection and assessment of tissue perfu-sion and vascular delivery of drugs or genes.202, 203 Ultrasoundcontrast imaging has succeeded in the detection of inflam-mation and angiogenesis of tumors, both of which play im-portant roles in tumor growth and metastasis.204–206 The gen-eral paradigm of ultrasound perfusion imaging is to acquire(i) a baseline (no contrast) image, (ii) subsequent dynamiccontrast-enhanced images following a bolus of microbubbles,and (iii) a high power ultrasound burst to rupture the mi-crobubbles and thus eliminate the contrast enhancement. Themicrobubbles are administrated intravenously and 1–2 ml ofcontrast agent is sufficient to produce good-quality imaging.Microbubbles manufactured to 2–4 μm diameters can pass

through the capillary bed but do not extravasate into the inter-stitium. Microbubbles may rupture spontaneously when theytravel under the ultrasound beam, depending on the level ofemitted mechanical energy. Their elimination is achieved viathe lungs following their trans-pulmonary passage and diffu-sion into the alveolar air. Microbubbles are completely elim-inated from the body approximately 20 min after intravenousinjection.207

Advantages that ultrasound offers are real-time imaging,noninvasiveness, nonionizing radiation, portability, low cost,and high spatial and temporal resolution. The resolution of ul-trasound molecular imaging ranges from 50 μm to 0.5 mm.This high spatial resolution is attainable by using frequenciesfrom 20 to 100 MHz. The primary competing characteristicsfor ultrasound molecular imaging are depth of penetration andspatial resolution, which vary inversely and directly, respec-tively, with increasing frequency.

Recently, targeted ultrasound imaging has been enhancedby developing new contrast agents such as liposomes andnanoparticles, which target particular intravascular receptorsor other structures.208, 209 Molecularly targeted microbubblesoffer the potential for personalized diagnostic imaging andtherapy through delivery of targeted contrast as well as inter-ventional agents.210 Use of receptor-specific ultrasound con-trast microbubbles to demonstrate highly vascularized tumorperiphery is illustrated in Fig. 11.

4. MOLECULAR IMAGING CHALLENGESIN CLINICAL RADIATION ONCOLOGY

4.A. Spatial scale in molecular imaging

The spatial scale of molecular imaging techniques coversapproximately 4 orders of magnitude, from the use of opti-cal and nanosensor techniques at the subcellular level (1 μmresolution) up to SPECT and MRSI (10 mm resolution).211

This wide variation in spatial scale presents challenges withrespect to integrating such data into a clinical radiation treat-ment planning system, such as molecular image registrationwith the corresponding planning CT. Figure 12 shows the spa-tial capabilities and relative importance for CT and molecu-lar techniques (see Ref. 212). A more detailed table of thelimiting spatial resolution for different imaging modalitiesis presented in Ref. 213. Although the spatial and tempo-ral resolution of various imaging modalities is improvingdue to technological advances, fundamental physical consid-erations in some cases may impose limits on the resolutionachievable.214–217

4.B. Image quality

Overall image quality depends on a number of complexinteracting factors, including the physical processes affectingthe signal and its origination (depth and surrounding tissues),acquisition sampling rates (spatial and temporal), backgroundsignal, target-to-background contrast, statistical uncertainty(or “noise”), scanner performance parameters, and the un-derlying biology which determines the distribution of the



FIG. 11. Ultrasound imaging of tumor vasculature by the use of ultrasound contrast microbubbles (MB) conjugated to the tripeptide sequence arginine-arginine-leucine (RRL). Nude mice bearing PC3 human prostate cancer cell lines were imaged 120 s after injection of either MBRRL or a MB control. Transverseultrasound image depicting persistent contrast enhancement in background subtracted images after injection of MBRRL [(a); white arrow], which is not seenin the control [(b); white arrow]. Noncolor coded portions are not background subtracted. The source of these images (From G. E. Weller et al., “Ultrasonicimaging of tumor angiogenesis using contrast microbubbles targeted via the tumor-binding peptide arginine-arginine-leucine,” Cancer Res. 65, 533–539 (2005).Copyright c© 2005 by American Association for Cancer Research.) states that the opacification is matching the anatomic pattern of the tumor vasculature, witha highly vascularized tumor periphery [white arrow in (a)] and a necrotic tumor core [white dashed arrow in (a)].

imaging signal in vivo. For instance, an ionizing photon emis-sion technique will suffer from signal degradation from inter-nal attenuation and scatter, and an optical tomography tech-nique will be dependent on photon scatter and sampling rate.Image quality of molecular images can be assessed using im-age analysis techniques developed for conventional imaging.Each molecular imaging modality is unique and will requirespecialized quality assurance and quality control.218, 219 Someimaging studies may require individual calibration or qual-ity assurance testing for each patient. Standardized phantoms,and QA tests and benchmark data for each imaging modalityfor a variety of lesion locations would be extremely valuablein helping users detect and/or avoid errors in image acquisi-tion and processing and application to therapy.

FIG. 12. Relative spatial ranges and strengths of molecular imaging tech-niques across modalities and spatial scale. Arrow height indicates relativefrequency of use (high or low). Arrow narrowing indicates decreased effec-tiveness of use as a function of spatial scale (From D. Piwnica-Worms, “Per-spectives in imaging: Present and future prospects of clinical molecular imag-ing,” in Holland-Frei Cancer Medicine, edited by D. Kufe, R. E. Pollock, R.R. Weiscselbaum, R. C. Bast, T. Gansler, J. F. Holland, and E. Frei (B.C.Decker, Inc., Hamilton, Ontario, 2003), pp. 551–556. Copyright c© 2003 byB.C. Decker Inc.).

4.C. Biologic structure definition and response

The use of molecular images in radiation treatmentpresents various challenges,220 including (1) image transmis-sion, (2) registration of multimodality images, (3) knowledgeof the information fidelity as well as image interpretation (tu-mor vs normal tissue vs inflammation) for the molecular im-age modality in use, and (4) composition of the target andcritical volumes from a set of multimodality, correlated imagesets. Issues of patient immobilization and motion, image fileformats, registration algorithms, and verification of softwaretools for processing and display also have to be addressed.

Problems associated with multimodality image registra-tion and fusion have been largely resolved for most clin-ical imaging modalities, including those with noisy (i.e.,sparse) data.16, 221–223 The introduction of multimodality,hybrid devices (e.g., PET-CT) has further simplified thisprocess.110, 224–228 However, molecular image sets may re-quire more robust registration algorithms and image dis-play and analysis capabilities in order to accommodate thewide range of spatial resolution, signal intensity, image ma-trices and depth, and voxel dimensions encountered acrosssuch modalities. Current commercial external-beam radiationtreatment planning systems likely cannot accommodate thevariety of molecular imaging data sets. The digital imagingand communications in medicine (DICOM) and DICOM - ra-diation therapy (DICOM-RT) image format standards will beimportant for “routine” use of molecular images.229, 230

Accurate image interpretation is required in order to usemolecular images for radiation oncology applications. Cor-rect image interpretation can yield information on the diag-nosis, classification (segmentation) of normal and target tis-sues, evaluation of a treatment plan, and treatment effective-ness. The seemingly simple task of setting the window-level



to manually draw a structure on a molecular image may mis-lead the user to draw a smaller or larger target volume thanwould be specified by using a quantitative (albeit a relative)index such as the SUV. Even target volume definition withFDG PET for lung cancer remains an elusive process withquite variable results depending on the approach used for de-termining the target boundary.231 Clearly, users of molecularimages will need to have experts who can interpret the bio-logical image and understand its clinical characteristics andlimitations. Robust software tools are also needed to aid ordirect the interpretation and delineation processes.

A fundamental understanding of the physical and techni-cal factors that affect the qualitative appearance and quanti-tative accuracy and precision of molecular images is neededto safely and meaningfully apply these images to therapy. Inother words, the biological significance must be known for theintensity level of a particular voxel for the molecular imagingtechnique. In vivo molecular imaging has been and will con-tinue to be used to assist in identifying cancerous tissues inthe body. The biological attributes of tumor and normal tis-sues cannot be assessed accurately if the information in themolecular images is biased. Common image artifacts affect-ing the accuracy of diagnostic PET images are routinely “readthrough” by the nuclear medicine physician or radiologist.However, if image artifacts were mistaken for areas of activetumor, they might be erroneously targeted for increased ra-diation dose, with potentially serious negative consequences.The accuracy and precision with which molecular imagingtechniques permit localization of regions for increased dosetargeting and response assessment are affected by the acqui-sition, image processing, and display methods.

Even with appropriate image manipulation and registrationtools and an understanding of the technical and biologicalcharacteristics of molecular images, the clinical use of suchimages is still challenged by the needs to define a target vol-ume. Current image-based radiation treatments use relativelysimple binary (or segmented) targets that define regions to betreated or avoided. With intensity modulated treatment, doseconstraints are used to optimize target volume coverage whilespecifying the dose limits for critical and normal tissue vol-umes. Although likely to be useful, biological imaging com-plicates the optimization process by (1) increasing the amountof potentially relevant information to enable “biological” con-straints, and (2) providing novel images about which the usermay be uncertain regarding image quality or interpretation.These considerations have been discussed in general17 andfor radiosurgery in particular.221 The reality is that the bio-logical target volumes for multimodality image sets will notbe congruent in size or shape. Temporal effects must also beaddressed when defining the target, an example of which isgiven by Sovik et al.11 In their study, the tissue oxygenationof canine sarcomas was measured by DCE-MRI throughoutthe course of fractionated radiotherapy to adaptively tailor theIMRT plan to tumor hypoxia. Although hypoxia imaging hasbeen one of the first uses of biologic imaging for clinical radi-ation oncology purposes, there has not been significant clin-ical success. Thus, the simple binary anatomical target be-comes one with a variable boundary which is biologically

matched to a radiation dose distribution of one or more doselevels for an individual patient.221 For instance, recent workshows that MR nerve tract diffusion tractography may pre-dict for pathways of spread for brain tumors, such that tractsoriginating in or near the gross tumor volume may requireinclusion in the radiation treatment field.232 A hypotheticalclinical target volume is shown, including the obvious diffu-sion tracks identified via MR diffusion imaging (Fig. 13). De-veloping a culture in which diagnostic radiologists, radiationoncologists, physicists, pathologists, basic scientists, vendors,and others routinely collaborate together to interpret, use, andperform quality assurance of molecular images is necessaryto determine the appropriate clinical applications for theseimages.

4.D. Biological modeling and application of molecularimaging to radiation oncology both for treatmentplanning and response assessment

Predictive models based on biological data from molec-ular images will provide information to the healthcare teamwhen making therapeutic decisions and prognoses. The re-cently published report of Task Group 166233 includes briefdescriptions and many references to current models for nor-mal tissue and tumor radiation response. Molecular imagesshould provide accurate spatial distributions of quantitativebiological data for tumor and normal-tissues, enabling the re-sulting biophysical models to be applicable to individual pa-tients. Image acquisition and processing techniques should bestandardized to insure image quality is acceptable for quanti-tative use in biological modeling of radiation dose response.

The principles of modern radiation oncology treatmentsunderscore the significance of image-based identification ofthe actual extent of viable tumors, to maximize tumor dosecoverage and uniformity and at the same time protectingsurrounding structures.23 Of particular interest, molecularimaging may be used in radiation oncology to augment bothtreatment planning and assessment of response to therapy.Modern high-resolution anatomical imaging combined withhigh-sensitivity molecular imaging can achieve volumetric tu-mor characterization and quantitative modeling of tissue irra-diation. In this way, more accurate planning of targeted andactual gross tumor volume may be achieved, while sparingneighboring sensitive organ structures from unnecessary ir-radiation. Noninvasive molecular imaging modalities, suchas PET and SPECT, are increasingly proposed for the col-lection of unique information about tumor radiobiology, forexample, hypoxia and glycolytic metabolism, and the useof such information allows for optimized “dose-painting” or“dose-sculpting” when implementing modern conformal orintensity-modulated radiation therapies.

5. INITIATIVES IN MOLECULAR IMAGING

Molecular imaging for cancer characterization, diagnosis,and treatment is a national priority. The National Cancer In-stitute (NCI), through its Cancer Imaging Program (CIP), isproviding numerous opportunities for research funding and



FIG. 13. Diffusion track imaging of potential pathways of spread for a primary brain tumor. (a) T1-weighted postcontrast image showing the primary anaplasticastrocytoma and treatment margin (pink line, 90% isodose) obtained from the treatment plan by computed tomography (b). (c) Diffusion tensor imaging referenceimage (b = 0) taken at the same time point showing all major diffusion paths passing through the tumor. Two major bundles pierce the treatment margin. (d)T1-weighted postcontrast image 3 months after the time point of (a) showing recurring tumor (green box) just outside of the treatment margin and in the samelocation as the major posterior bundle (c). (e) The treatment plan that was used for stereotactic radiotherapy (pink line) and the proposed anisotropic treatmentplan (green line) with increased dose along the two prominent bundles emanating from the primary tumor and reduced margin along other directions. (f)Depiction of the original (pink) and proposed (green) treatment plans of the follow-up T1-weighted postcontrast image. Their contention was that the proposedplan might have stalled or may have even prevented the onset of the secondary tumor. Reprinted from A. P. Krishnan, I. M. Asher, D. Davis, P. Okunieff, and W.G. O’Dell, “Evidence that MR diffusion tensor imaging (tractography) predicts the natural history of regional progression in patients irradiated conformally forprimary brain tumors,” Int. J. Radiat. Oncol., Biol., Phys. 71, 1553–1562 (2008). Copyright c© 2008 by Elsevier Inc..

policy setting in the area of novel cancer imaging with themission to promote and support “cancer-related research inimaging sciences and technology for the understanding ofcancer biology and . . . the management of cancer and can-cer risk.”234 The NCI-driven quantitative imaging initiativerequires substantial medical physics input to be successful.Opportunities exist at all levels from the basic to clinicalsciences and for academic and commercial entities withinthe following programs: (1) American College of RadiologyImaging Network, (2) Network for Translational Research:Optical Imaging, (3) Development of Preclinical Drugs andEnhancers, (4) In Vivo Cellular and Molecular Imaging Cen-ters, (5) Interagency Council on Biomedical Imaging in On-cology, and (6) Small Animal Imaging Resource Program.The National Institute for Biomedical Imaging and Bioengi-neering (NIBIB) has several programs for research fundingin the areas of development of novel drug and gene deliv-ery, image-guided interventions, improvements in imagingmethods, systems and methods for cellular and molecularimaging, systems and methods for small animal imaging, nu-clear medicine, optical imaging and spectroscopy, ultrasound,and x-ray, electron, and ion beam imaging.235 Other oppor-

tunities for education and collaboration have been affordedby the Biomedical Imaging Research Opportunities Work-shops (BIROW)236 and by symposia on molecular and on-cologic imaging being hosted by scientific societies such asthe AAPM, the RSNA, the American Society of RadiationOncology (ASTRO), and the American College of Radiology(ACR).

6. CONCLUSIONS

Molecular imaging is the imaging of specific, molecularsources using radioisotope, magnetic resonance, optical, orultrasound methods. Oncologic molecular imaging has beenestablished as a national priority by government and scientificorganizations, and holds great promise for patients throughimaging of tumor and normal-tissue environments to providein vivo characterization of hypoxia, perfusion, apoptosis, andnutritional state. For the clinical application of these tech-niques, spatial scale issues related to accurate registration andclinical interpretation of data over 4 orders of magnitude needto be resolved. Such an approach is a paradigm shift for thefield of radiation oncology. Molecular and biological targets



as well as volumes from digital images may be used to biolog-ically optimize radiation dose distributions. Image informa-tion in spatial, temporal, physical, and biological intensity for-mats may be used quantitatively. Radiation treatment could beaccomplished using “conventional” doses delivered with in-tensity modulated approaches or through lower doses to yieldradiation-initiated molecular events.3 In particular, PET imag-ing can readily be found in radiation oncology clinics.

Molecular imaging and advanced imaging in radiation on-cology present great challenges and opportunities for col-laborations through the convergence of molecular biology,diagnostic radiology, radiation oncology, physics, imagingscience, chemistry, and other fields. Integration of diagnos-tic radiology and radiation oncology, in particular, is under-way, with imaging providing the key link among diagnosis,treatment, and response assessment.3 A growing communityof multidisciplinary scientists and clinicians is required toaddress the acquisition, manipulation, interpretation, and ap-plication of molecular and biological images in the clinicalsettings.

ACKNOWLEDGMENTS

The authors wish to thank the reviewers from both theAAPM’s Therapy Imaging Subcommittee and the TherapyPhysics Committee for their thoughtful and helpful com-ments. The authors also wish to thank Dr. Akira Kawashima,Dr. Ken Wheeler, Dr. Bradley Kemp, and General ElectricHealthcare for some of the figures shown in this paper.

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