DOI: 10.1148/radiol.2381041602
(Radiology 2006;238:16-39.)
© RSNA, 2006
Evolving and Experimental Technologies in Medical Imaging1
Anthony B. Wolbarst, PhD and
William R. Hendee, PhD
1 From the Department of Radiation Medicine, Georgetown University Medical School, Washington, DC (A.B.W.); and Medical College of Wisconsin, 8701 Watertown Plank Rd, Milwaukee, WI 53226 (W.R.H.). Received September 20, 2004; revision requested November 15; revision received December 9; accepted January 14, 2005; updated July 14; final version accepted August 5. Supported in part by grants from the National Institutes of Health (RO1 CA80490, P01 CA 87634) and GE Medical Systems.
Address correspondence to W.R.H. (e-mail: whendee{at}mcw.edu).
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ABSTRACT
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Medical images are created by detecting radiation probes transmitted through or emitted or scattered by the body. The radiation, modulated through interactions with tissues, yields patterns that provide anatomic and/or physiologic information. X-rays, gamma rays, radiofrequency signals, and ultrasound waves are the standard probes, but others like visible and infrared light, microwaves, terahertz rays, and intrinsic and applied electric and magnetic fields are being explored. Some of the younger technologies, such as molecular imaging, may enhance existing imaging modalities; however, they also, in combination with nanotechnology, biotechnology, bioinformatics, and new forms of computational hardware and software, may well lead to novel approaches to clinical imaging. This review provides a brief overview of the current state of image-based diagnostic medicine and offers comments on the directions in which some of its subfields may be heading.
© RSNA, 2006
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INTRODUCTION
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Medical images are produced through a variety of processes that make use of physical probes that are created, affected by the body, and detected in strikingly different ways. Photons of all energies (x-ray, gamma ray, annihilation, ultraviolet, optical, infrared, microwave, radiofrequency), weak electric and magnetic fields, ultrasound waves, and other probes differ substantially in their ability to penetrate the body, in the types of noise with which they must compete, and in the ease with which they can be detected and localized. These factors, along with concerns related to radiation dose, acoustic power, very strong static and dynamic magnetic fields, et cetera, can influence the level of contrast achievable among healthy and diseased tissues, the spatial and temporal resolutions possible, the presence of noise or artifacts, and the overall clinical utility of an imaging modality.
As a result, there are a number of widely accepted standard film-based and digital technologies to choose from in posing and addressing a particular clinical question. Part of the task of the clinician is to be adept at selecting the one with the greatest likelihood of providing the correct diagnostic answer safely and at acceptable cost.
There are also several nonstandard imaging modalities, some new and others that have been around for awhile. These are being explored and may (or may not) eventually find routine clinical application. A few of them, such as electroencephalography, magnetocardiography, and thermography, create images from extremely faint signals produced by the body itself. Tissue impedance tomography and diaphanography operate, like x-rays, on the basis of observations of how probes passing through the body interact with it.
Meanwhile, biomedical research is becoming ever more interdisciplinary, as it depends increasingly on cross fertilization between the physical sciences and engineering on the one hand and biology and clinical medicine on the other. Examples of the fruits of such collaboration abound—such as recent developments in molecular imaging, biomedical informatics, nanobiotechnology, computer-aided detection and diagnosis, and the merging of diagnosis and therapy as in image-guided intervention.
This review will provide a brief overview of the current state of image-based diagnostic medicine and offer comments on the demonstrated or potential clinical efficacy of some of the newer imaging modalities. Additional viewpoints may be found in the reviews and other articles listed among the references.
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ADVANCES IN CURRENT TECHNOLOGIES
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The Table provides representative values for some of the imaging parameters of typical state-of-the-art commercial image-acquisition devices; the entries for any particular model may, of course, differ considerably from these representative values.
Planar X-ray Imaging
There are two general categories of imaging with the high-energy photons produced by x-ray tubes: (a) planar-projection approaches, such as standard radiography (whether film-based or digital) and fluoroscopy (increasingly with a fully digital image receptor) and (b) tomographic approaches, such as CT. Either way, image information is obtained initially in the form of shadows cast in an x-ray beam through differential attenuation or phase shifting (1) by materials in the body that differ in thickness, density, or chemical makeup (ie, atomic number). There have been giant strides forward, over the past few years, in both categories.
In planar x-ray imaging, new electronic image receptors are taking over the field (2). The standard combinations of fluorescent screen plus x-ray film and of image intensifier plus television monitor or charge-coupled device optical camera have served radiology well for a good part of a century, but their evolution has largely plateaued. Meanwhile, digital image receptors keep on getting better, and fully digital systems, built around picture archiving and communication systems (PACS), may soon be dominating x-ray projection imaging. A number of radiology departments already are filmless, and the direction of the trend is clear (3). Even in breast imaging, where resolution is essential for the assessment of microcalcifications, digital systems are becoming increasingly accepted (4,5).
Photostimulable phosphor plates composed of BaFBr and BaFI have been used in computed radiography cassettes since the 1970s. More recently, the real-time, active-matrix, flat-panel imager of digital radiology and digital fluoroscopy has begun displacing its analog predecessors. A flat-panel imager is an array of hundreds, thousands, or millions of tiny independent semiconductor detectors that are themselves sensitive either to high-energy photons (direct-detection array) or to light from an adjacent thin layer of fluorescent material (indirect-detection array) (6,7). It is possible that recently developed ink-jet printing (rather than photographic) lithographic technology will allow the fabrication of flat-panel imagers on flexible plastic substrates of virtually any size.
There are advocates of both computed and digital radiography, but the two are complementary modalities, and there is room for both (8). Computed radiography cassettes are somewhat more flexible—coming in a range of sizes and shapes—and more cost-effective: To install, one need only exchange a film cassette assembly with a computed radiography cassette; also, a damaged computed radiography cassette can be replaced at reasonable cost. Digital radiography allows faster throughput and is more dose effective, but it is also more expensive to implement. Both are likely to be around for a while, providing quality images of teeth, the chest, bone, the gastrointestinal tract, the extremities, and the breast (9,10) and contributing to a growing variety of interventional procedures.
Among the great advantages of both technologies (and of digital imaging in general) are (a) the linearity of response of digital radiation receptors over exposure ranges that are orders of magnitude greater than the latitude of any film and the overall improvement in image quality that results from that property; (b) the separation of image acquisition from image processing (windowing, edge enhancement, noise reduction, etc) and from display, so that these three steps can be optimized independently; and (c) the overall benefits of digital signal storage, communication, and analysis.
Through the digital subtraction of images created with different x-ray tube kilovoltage settings, for example, dual- (and higher-) energy methods (11) can be used to remove unwanted visual background noise, such as bonelike or soft-tissue structures, that obscures what the clinician is looking for. In a related vein, dual-energy x-ray absorptiometry is the standard method for measurement of bone density, the major predictor of fracture due to osteoporosis. However, dual-energy x-ray absorptiometry is now being challenged by US and other techniques (12,13).
Digital tomosynthesis takes a different approach to achieving the same end (14,15). Like its film-based predecessor, digital tomosynthesis generates multiple planar images at arbitrary depths within a patient. A series of discrete projection radiographs are acquired with digital detectors while the x-ray tube and beam are pivoted through a small angle. Through manipulation of the projection images to remove the over- and underlying planes, high-quality images of each selected tomographic plane can be obtained with section thicknesses ranging from 1 mm to several centimeters (Fig 1). Tomosynthesis is a type of limited-angle tomography and has relatively poor resolution in the depth direction. The level of detail in the imaging plane, however, can be superb. Digital tomosynthesis has been applied to chest, breast, angiographic, orthopedic, dental, and other kinds of examinations.
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Figure 1a: Digital tomosynthesis of human lung. (a) Posteroanterior radiograph shows region of interest. (b) Digital tomosynthesis section demonstrates 15-mm pulmonary nodule (arrow) that was not visualized in a. Note also improved clarity of vascular detail in b, which was reconstructed with a matrix inversion tomosynthesis technique that used 61 projection images acquired in 10 seconds over a total tube swing angle of 16°. Fifty-nine sections were generated with 3-mm spacing. Total subject entrance exposure was approximately the same as that for a screen-film lateral chest radiograph. (Image courtesy of James T. Dobbins III, PhD, H. Page McAdams, MD, and Devon J. Godfrey, Duke University Medical Center.)
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Figure 1b: Digital tomosynthesis of human lung. (a) Posteroanterior radiograph shows region of interest. (b) Digital tomosynthesis section demonstrates 15-mm pulmonary nodule (arrow) that was not visualized in a. Note also improved clarity of vascular detail in b, which was reconstructed with a matrix inversion tomosynthesis technique that used 61 projection images acquired in 10 seconds over a total tube swing angle of 16°. Fifty-nine sections were generated with 3-mm spacing. Total subject entrance exposure was approximately the same as that for a screen-film lateral chest radiograph. (Image courtesy of James T. Dobbins III, PhD, H. Page McAdams, MD, and Devon J. Godfrey, Duke University Medical Center.)
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Very fast digital technology makes scatter rejection possible through time-of-flight methods. On the other hand, it is possible to create pictures from scattered (rather than transmitted) radiation, as is done with some security-checking devices (Fig 2). It even appears that information gleaned from tissue-induced phase shifts (which are quantified in terms of wave optics, rather than the ray optics of conventional x-ray imaging) may lead to improved soft-tissue contrast at no increase in patient exposure.
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Figure 2: Dependence of relative amplitude of x-ray scatter (in this case, through 180°) on atomic number. Truck is being irradiated from the side, and backscattered radiation is imaged by an image receptor on the same side as the radiation source. Organic and other materials with a lower atomic number, such as drugs, explosives, plastic weapons, or people, scatter x-rays much more readily than do high-atomic-number items, such as trucks and guns. The amount of back-scatter from an object also depends, of course, on its physical density. (Image courtesy of American Science and Engineering, Billerica, Mass.)
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CT Imaging
As with contrast agent or dual-energy digital subtraction techniques, and as with digital tomography, the power of CT lies in its ability to remove irrelevant, unwanted, and interfering information. Just as a radiologist may be able to determine the shape of an object in a patient by viewing radiographs taken from several angles, so also does CT involve acquisition of a large number of x-ray projection views. CT, in effect, works backward to reconstruct the spatial distribution of materials—or, more precisely, the distribution of their x-ray photon attenuation characteristics—that gave rise to those projections. Two fairly recent developments have substantially increased the efficiency and speed of data acquisition in CT (16).
In axial-mode nonhelical CT systems, the tube is rotated once about an immobile table; the table is then advanced by a small amount in the z-axis direction, and the tube is rotated once in the opposite direction. This procedure is repeated over and over, with each rotation producing data for a flat plane of tissue. With helical CT, also known as spiral CT, the table and x-ray tube both move continuously throughout data acquisition. The technology is either so-called third generation, with a rotating array of detectors, or fourth generation, in which a thousand or so very small detectors surround the patient and are fixed in space. Slip rings are needed to deliver electric power to the tube and, for third-generation machines, to retrieve the signals from the detectors (Fig 3). Helical CT allows much faster scanning and provides images with higher resolution in the transverse plane.
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Figure 3a: Slip rings used to bring power to x-ray tube on rotating gantry of a helical CT machine and, for some designs, to acquire information from the detector array. (a) The shiny metal strips carry electric signals that are swept off by special brushes. (b) The brushes are not in the form of bristles but rather of metal blocks (in this case a silver alloy). The five pairs of larger brushes provide the voltage required by the x-ray tube, and the three pairs of smaller ones transfer signals from the gantry controller.
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Figure 3b: Slip rings used to bring power to x-ray tube on rotating gantry of a helical CT machine and, for some designs, to acquire information from the detector array. (a) The shiny metal strips carry electric signals that are swept off by special brushes. (b) The brushes are not in the form of bristles but rather of metal blocks (in this case a silver alloy). The five pairs of larger brushes provide the voltage required by the x-ray tube, and the three pairs of smaller ones transfer signals from the gantry controller.
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The second important innovation is the transition from single- to multisection imaging (17,18). The single, narrow, transverse fan beam of x-rays used for axial mode CT is broadened into a cone beam, and each detector is replaced with a row of four to 64 (at present) separate detectors aligned parallel to the axis of the patient—like a belt with four to 64 thin rows of small closely spaced studs running along its length. This approach to efficiency in data acquisition, long incorporated into PET, makes it possible for a multi–detector row scanner (also known as a multidetector or multisection scanner) to acquire images of up to 64 sections of the body simultaneously. This yields a dramatic improvement in three-dimensional display quality, since small steps between sections lead to smooth, high-resolution, three-dimensional images. Three-dimensional CT angiography, for example, is an increasingly important, and much less invasive and costly, alternative to standard catheter-directed coronary arteriography for the demonstration of coronary artery disease, the detection and measurement of the volume of noncalcified plaque, and other applications.
Multi–detector row CT scanners use classic third-generation scan geometry but with a cone- rather than a fan-shaped x-ray beam. For 16 or fewer sections, it is possible to employ well-established multi–parallel-section approximate reconstruction algorithms. More sophisticated approximate and exact cone-beam algorithms have been created and yield better images, but they are computationally more complex and time consuming (19–22).
CT scanner manufacturers continue to strive for greater scanning speed and more longitudinal anatomic coverage with each gantry rotation. Because the centrifugal forces would become overwhelming, the speed of gantry rotation will probably not increase much beyond its present value of three cycles per second. This is good enough to examine the lung (23,24) or capture a heart at systole or diastole with little motion blurring, but shorter scan times are clinically desirable and might be achievable with multiple tubes. So, too, is z-axis coverage, to allow functional imaging of the heart, lung, brain, and other structures. These, and the development of better detector systems, will further improve resolution in the transverse dimension, so that CT is becoming a truly isotropic modality, similar to MR imaging. A difficulty with this, of course, is the increase in Compton scatter radiation reaching the image receptors. Indeed, one of the touted benefits of single-section scanning, in the early days, was the virtual absence of scatter—and the best approach to dealing with it, now that it has returned with multisection imaging, is not obvious (25,26).
Researchers at the University of Aachen, Germany, have reported on a device in which the multiple rings of detectors have been replaced with a flat-panel imaging array (26,27). Several manufacturers have already produced flat-panel prototypes, and quasi-CT C-arm devices have been marketed. However, there are daunting problems associated with the very high rates (gigabytes per second) of data transfer.
With dual- or multienergy techniques, moreover, CT could provide information on tissue parameters beyond simple pixel-average relative attenuation coefficients (ie, Hounsfield units).
Revolutionary when it was introduced in the mid-1980s, electron-beam CT can produce a transverse section in as short a time as 50 msec, fast enough to freeze the beating of the heart in any phase of the cardiac cycle, and at a relatively low dose—about 1 mSv for angiography, compared with 10 times that with 64-section CT—and even less for coronary calcium scoring (Fig 4). Electron-beam CT images tend to be noisy, however, and the far more versatile 64-section CT has better resolution in all three directions (about 0.3 vs 1.5 mm for electron-beam CT). Fast MR imaging can also support cardiac screening or image-guided intervention and can do so without ionizing radiation.
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Figure 4a: Electron-beam CT, also known as fifth-generation CT. (a) Diagram shows electron-beam CT scanner architecture. Target is a long continuous tungsten strip that makes a 210° arc around the patient. Electron beam and focal spot traverse the entire arc in 50–100 msec. Electron beam and target reside in a single funnel-shaped evacuated chamber (not shown). Detector array does not move. (b) Transverse electron-beam CT image in a patient with obstructive coronary artery disease who had undergone bypass surgery. Calcium is visible in left main coronary artery (lower left arrow), left anterior descending coronary artery (lower right arrow), and aorta (upper left arrow). Surgical clip (upper right arrow) from bypass surgery is also visible.
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Figure 4b: Electron-beam CT, also known as fifth-generation CT. (a) Diagram shows electron-beam CT scanner architecture. Target is a long continuous tungsten strip that makes a 210° arc around the patient. Electron beam and focal spot traverse the entire arc in 50–100 msec. Electron beam and target reside in a single funnel-shaped evacuated chamber (not shown). Detector array does not move. (b) Transverse electron-beam CT image in a patient with obstructive coronary artery disease who had undergone bypass surgery. Calcium is visible in left main coronary artery (lower left arrow), left anterior descending coronary artery (lower right arrow), and aorta (upper left arrow). Surgical clip (upper right arrow) from bypass surgery is also visible.
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Meanwhile, a novel design of a standard x-ray tube is being tested clinically. This system allows anode cooling that is five or so times faster than is possible with a standard rotating anode tube, making feasible either correspondingly longer uninterrupted exposure durations or more intense (hence, briefer) pulses (28).
While biomedical informatics is perhaps best known for its role in untangling the vast quantities of data coming from work on the genome, another area of considerable interest is the need to display the information content of hundreds, perhaps even thousands, of CT sections per patient (with numerous such patients each day) in diagnostically useful ways. A radiologist clearly cannot examine them all one by one (despite the possibility that a critical clinical sign may appear faintly in one of them—with serious legal implications). Conversely, the use of a helical 64-section machine just to create a smoother surface rendering of some organ would rarely be justified—although in some areas that require very high resolution, such as CT mammography or CT virtual colonoscopy, it might be justified (29). So, it is essential to develop computer-based bioinformatics programs that can locate, and perhaps interpret, the few clinically essential visual items contained within such an overabundance of imaging data (30).
Finally, there is strong pressure to reverse the upward "radiation dose creep" for all x-ray technologies, especially for CT in children and for mass screening applications. Fortunately, most operators of the machines appear to be increasingly aware of the need to tailor the technique factors to the dimensions of the patient, which can reduce the dose dramatically (31,32).
Gamma Ray Imaging
In standard nuclear medicine, images are produced by using radiopharmaceuticals and a gamma camera (33). A radiopharmaceutical consists of two parts: (a) an agent designed to concentrate, through chemical or other physiologic means, in a specific organ or compartment of the body and (b) a radioisotope (most commonly, technetium 99m), which emits a medium-energy gamma photon. While a nuclear imaging study cannot compete with radiography or CT in revealing precise anatomic detail, it may more than compensate by providing potentially invaluable information about the physiologic status of an organ or other tissue.
Radiopharmacology continues to evolve rapidly, and it is to be expected that scintillation-detector materials for gamma image receptors will continue to be improved and that, at some point, either the photomultiplier tubes or the entire scintillation crystal–photomultiplier tube assembly will be replaced by sensitive, low-noise, solid-state detectors, perhaps of the flat-panel imager variety.
SPECT, the gamma ray counterpart to CT, has become nearly indispensable for the assessment of myocardial perfusion and other cardiac functions. More than half of SPECT studies are for coronary artery disease, another quarter are for bone, and the rest are for brain, prostate, thyroid, and other organs (34,35).
Likewise, PET has long enjoyed a prominent role in neurologic research. More recently, PET has been used as a primary clinical tool for the detection, localization, staging, and monitoring of malignancies (34,36–38). PET makes use of radionuclei that emit positrons, the positively charged antiparticles of electrons. Positron emitters include oxygen 15, nitrogen 13 (13N), carbon 11, rubidium 82, and, the most extensively used one fluorine 18 (18F), which is normally employed in the form of the glucose analog fluorodeoxyglucose. A positron travels a few millimeters in tissue and collides with an atomic electron. The particles annihilate one another and create a pair of 511-keV "annihilation" photons that fly off in almost exactly opposite directions. Only events that trigger two of the PET device's detectors on opposite sides of the patient and in coincidence (nearly simultaneously) contribute to the creation of the PET image. The operation of opposed detectors in coincidence improves both the spatial resolution and the signal-to-noise ratio of the images.
The number of clinical PET and fusion PET/CT procedures in the United States has grown from about 250 000 in 2001 to roughly 900 000 in 2004. Ninety percent of the studies are performed in search of tumors; the remainder are split between cardiac and neurologic studies. It is possible that, because of its better resolution, physiologically more interesting agents, and decreasing costs, PET will come to displace SPECT and conventional nuclear medicine for many of the latter modalities' current roles.
The fusion of SPECT or PET images with CT or MR images (39,40) allows the projection of visual physiologic information on a detailed anatomic map background so that tissue function can be correlated directly with tissue structure (Fig 5). With fusion, moreover, the CT attenuation data also make possible substantial correction of the PET or SPECT image for attenuation of gamma rays within the body. It has even proved useful in radiation treatment planning to merge CT and MR images (Fig 6).
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Figure 5: PET/CT fusion imaging. Coronal CT (left), PET (middle), and fused PET/CT (right) images show distribution of positron-emitting 18F-fluorodeoxyglucose superimposed on CT display of anatomy. (Image courtesy of Robert Hellman, MD, Medical College of Wisconsin, Milwaukee, Wis.)
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Figure 6a: Transverse MR/CT fusion imaging for radiation therapy treatment planning. Brain tumor is barely visible on (a) CT scan but is clearly evident on (b) MR image (repetition time msec/echo time msec, 10 000/138) and (c) fused MR/CT image. At present, nearly all treatment-planning systems used to generate isodose maps for radiation therapy can work with CT or with fused MR/CT images (as in d) but not with MR images alone. This situation is likely to improve soon. (Image courtesy of Allen Li, PhD, Medical College of Wisconsin, Milwaukee, Wis.)
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Figure 6b: Transverse MR/CT fusion imaging for radiation therapy treatment planning. Brain tumor is barely visible on (a) CT scan but is clearly evident on (b) MR image (repetition time msec/echo time msec, 10 000/138) and (c) fused MR/CT image. At present, nearly all treatment-planning systems used to generate isodose maps for radiation therapy can work with CT or with fused MR/CT images (as in d) but not with MR images alone. This situation is likely to improve soon. (Image courtesy of Allen Li, PhD, Medical College of Wisconsin, Milwaukee, Wis.)
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Figure 6c: Transverse MR/CT fusion imaging for radiation therapy treatment planning. Brain tumor is barely visible on (a) CT scan but is clearly evident on (b) MR image (repetition time msec/echo time msec, 10 000/138) and (c) fused MR/CT image. At present, nearly all treatment-planning systems used to generate isodose maps for radiation therapy can work with CT or with fused MR/CT images (as in d) but not with MR images alone. This situation is likely to improve soon. (Image courtesy of Allen Li, PhD, Medical College of Wisconsin, Milwaukee, Wis.)
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Figure 6d: Transverse MR/CT fusion imaging for radiation therapy treatment planning. Brain tumor is barely visible on (a) CT scan but is clearly evident on (b) MR image (repetition time msec/echo time msec, 10 000/138) and (c) fused MR/CT image. At present, nearly all treatment-planning systems used to generate isodose maps for radiation therapy can work with CT or with fused MR/CT images (as in d) but not with MR images alone. This situation is likely to improve soon. (Image courtesy of Allen Li, PhD, Medical College of Wisconsin, Milwaukee, Wis.)
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Fused images can be acquired with a single composite hybrid machine (eg, CT/PET scanner) that acquires and integrates the two separate imaging studies one immediately after the other, with the patient on the same table. Alternatively, one can combine and register the data from two independent devices with specialized software, which is far more demanding with the body than with the head and is subject to distortions due to changes in patient positioning. Common platforms for acquisition and display of fused image data from different imaging methods will undoubtedly continue to develop as image-guided surgery and radiation therapy evolve.
Meanwhile, advances in molecular biology are yielding receptor-specific radiopharmaceuticals and other pharmaceuticals that are opening up a new realm of opportunities for in vivo imaging at the molecular and cellular scale. Research in radiopharmacology and in the technologies for producing radionuclides are areas of importance but of uncertain near-term funding.
MR Imaging
A hydrogen nucleus (proton) in a water or lipid molecule in a cell acts rather like a spinning positively charged ball. As with any other moving charged body, it produces its own local magnetic field, somewhat like that of a compass needle.
When placed in a strong (eg, 1.5- or 3-T) external magnetic field, a proton tends to settle down comfortably into its ground state, also like a compass needle, with its spin axis and local field aligned along that of the external field. But a proton follows the dictates of quantum mechanics, not of common sense, and it can also reside temporarily in its metastable, higher-energy spin state, pointing in the "wrong" direction. Indeed, a group of protons can be intentionally flipped or twisted upside down and into their higher-energy spin states in a process known as nuclear magnetic resonance through the absorption of a pulse of radiofrequency energy of the correct (ie, Larmor) frequency and duration.
The spins subsequently experience naturally occurring Larmor frequency magnetic noise, which tickles them back down to their lower-energy configuration. This relaxation occurs rapidly if the water molecules are tumbling and interacting with one another and with their environments in such a manner that, at least intermittently, they are rotating at the Larmor frequency. In this condition, each proton experiences the magnetic field from its partner as varying, at least briefly, at the Larmor frequency, which causes a downward spin transition. The average spin-relaxation time for this transition is called T1. Free water tumbles much faster than the proton Larmor frequencies for the external fields employed in MR imaging, so T1 is long (4000 msec or so); but for water bound to intermediate-sized biologic molecules (which themselves may happen to be tumbling at the Larmor frequency), T1 can be an order of magnitude shorter.
MR imaging involves manipulation of pulses of radiofrequency energy and magnetic field gradients to assess, in effect, T1 and the related parameter T2 at each point in the part of the body being examined (41,42). MR imaging is of diagnostic value because it can be used to generate two- or three-dimensional maps that reflect the spatial distribution of the values of T1 and T2. But the proton spin-relaxation times depend largely on the degree of binding of the water molecules (in which the proton spin flips are occurring) to the nearby biologic molecules, and the water–biologic molecule interactions in a tissue are sensitive to its histologic characteristics and also to its physiologic status. So a T1- or a T2-weighted MR image can reveal information on both the anatomy and the state of health of tissues. Similar considerations apply for hydrogen atoms in lipids.
T1 and T2 can be altered in clinically useful ways with MR contrast agents, many of which are built around paramagnetic gadolinium ions caged within molecules of a chelating agent (43). Dynamic contrast enhancement and related techniques, where the proton spin-relaxation times in a region are monitored during the administration of a bolus of contrast agent, can reveal much about the microvascular environment.
MR imaging will continue to find new areas in which to expand (44). It has already spun off MR angiography (a competitor with CT angiography) (45–48) and MR microscopy (49). Also, single-wire or other radiofrequency antennae that can be positioned within vessels, body cavities, and elsewhere may provide gateways to new types of studies of the heart and other accessible organs. Diffusion MR imaging, which is becoming more widely used in the assessment of acute stroke and malignancies, is able to reveal information (the diffusion tensor) about the actual directions of blood diffusion (50,51).
MR spectroscopy and the closely related modality of chemical shift MR imaging can elucidate chemical processes ongoing in small volumes of tissue (52). As with fusion PET/CT, MR spectroscopy is being combined with high-resolution MR imaging to enhance sensitivity and specificity in the detection and identification of cancers and other abnormalities (Fig 7). And MR spectroscopy of carbon 13, fluorine 19, and phosphorus 31 (rather than hydrogen 1) is also providing valuable insights on cellular metabolism and biochemistry (53).
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Figure 7: Chemical shift MR/MR spectroscopic imaging (1000/18; bandwidth, 1000 Hz; 512 points with two signals acquired). Transverse short-echo-time chemical shift image (left) acquired at 0.5 T in a presymptomatic patient with Huntington disease shows strong elevation of glutamate (right: upper spectrum) in the head of the putamen. Unaffected thalamus (right: lower spectrum) is shown for comparison. (Image courtesy of Robert Prost, PhD, Medical College of Wisconsin, Milwaukee, Wis.)
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There is also movement toward devices for special applications, such as machines dedicated exclusively to breast imaging. Breast MR imaging (54,55) is especially promising for women who are at high risk for breast cancer or who have unusually dense breast tissue, which presents a severe challenge to screen-film and digital mammography (Fig 8). Likewise, some surgical suites are being designed so that an MR magnet can readily be wheeled into the operating room to ensure complete removal of a tumor at the time of surgery; alternatively, patient tables can be moved briefly from any of several operating rooms to a central magnet.
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Figure 8a: Mammography and MR imaging of the breast. (a) Left craniocaudal and (b) left mediolateral oblique mammograms demonstrate an irregular high-density mass (*). Overall, breast tissue density is heterogeneous, which is consistent with patient's age (35 years). Biopsy results showed the mass to be grade III invasive ductal carcinoma. (c, d) Transverse three-dimensional fast spoiled gradient-echo MR imaging (21.4/4.2) was performed to assess remaining breast tissue. (c) Enhancing mass corresponds to the known biopsy-proved breast cancer. (d) Incidental mass (arrow) was found posteriorly in the same breast. Biopsy was performed later with US guidance, and the mass found to be grade I invasive ductal carcinoma. This mass was not evident on the mammogram, either initially or in retrospect. (Image courtesy of Lonie Salkowski, MD, Medical College of Wisconsin, Milwaukee, Wis.)
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Figure 8b: Mammography and MR imaging of the breast. (a) Left craniocaudal and (b) left mediolateral oblique mammograms demonstrate an irregular high-density mass (*). Overall, breast tissue density is heterogeneous, which is consistent with patient's age (35 years). Biopsy results showed the mass to be grade III invasive ductal carcinoma. (c, d) Transverse three-dimensional fast spoiled gradient-echo MR imaging (21.4/4.2) was performed to assess remaining breast tissue. (c) Enhancing mass corresponds to the known biopsy-proved breast cancer. (d) Incidental mass (arrow) was found posteriorly in the same breast. Biopsy was performed later with US guidance, and the mass found to be grade I invasive ductal carcinoma. This mass was not evident on the mammogram, either initially or in retrospect. (Image courtesy of Lonie Salkowski, MD, Medical College of Wisconsin, Milwaukee, Wis.)
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Figure 8c: Mammography and MR imaging of the breast. (a) Left craniocaudal and (b) left mediolateral oblique mammograms demonstrate an irregular high-density mass (*). Overall, breast tissue density is heterogeneous, which is consistent with patient's age (35 years). Biopsy results showed the mass to be grade III invasive ductal carcinoma. (c, d) Transverse three-dimensional fast spoiled gradient-echo MR imaging (21.4/4.2) was performed to assess remaining breast tissue. (c) Enhancing mass corresponds to the known biopsy-proved breast cancer. (d) Incidental mass (arrow) was found posteriorly in the same breast. Biopsy was performed later with US guidance, and the mass found to be grade I invasive ductal carcinoma. This mass was not evident on the mammogram, either initially or in retrospect. (Image courtesy of Lonie Salkowski, MD, Medical College of Wisconsin, Milwaukee, Wis.)
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Figure 8d: Mammography and MR imaging of the breast. (a) Left craniocaudal and (b) left mediolateral oblique mammograms demonstrate an irregular high-density mass (*). Overall, breast tissue density is heterogeneous, which is consistent with patient's age (35 years). Biopsy results showed the mass to be grade III invasive ductal carcinoma. (c, d) Transverse three-dimensional fast spoiled gradient-echo MR imaging (21.4/4.2) was performed to assess remaining breast tissue. (c) Enhancing mass corresponds to the known biopsy-proved breast cancer. (d) Incidental mass (arrow) was found posteriorly in the same breast. Biopsy was performed later with US guidance, and the mass found to be grade I invasive ductal carcinoma. This mass was not evident on the mammogram, either initially or in retrospect. (Image courtesy of Lonie Salkowski, MD, Medical College of Wisconsin, Milwaukee, Wis.)
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A major objective of MR researchers and manufacturers is to reduce image-acquisition time without loss of image quality (56,57). Some are designing special radiofrequency and gradient pulse sequences that are now fast enough to capture the beating of a heart in cine form. Others have developed methods that involve energizing and obtaining signals from multiple radiofrequency coils simultaneously, a process known as parallel imaging, which substantially reduces the number of phase-encoding pulses required and, therefore, the imaging time (58).
Functional MR imaging, in particular blood oxygen level–dependent imaging, makes use of fast pulses to determine where the flow of oxygenated blood is unusually high and, by inference, where neurons are unusually busy. This method provides information on neural activity that is complementary to contributions from MR spectroscopy and PET (59,60). In a study of potentially staggering implications, functional MR imaging allowed researchers to determine which of several visual patterns that subjects were looking at—in effect, it allowed researchers to read the subjects' mind (61–63).
The nuclear MR signals that underlie MR imaging can be influenced by the magnetic coupling of nuclear spins that are far apart—even millimeters apart. The possibility of such an interaction, previously ignored because of its extreme weakness, has now given birth to a new branch of MR known as zero-quantum imaging (64). Still in the early developmental stage, zero-quantum imaging may be unusually sensitive for the detection of malignant tumors (Fig 9).
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Figure 9a: Transverse intermolecular double-quantum coherence (IDQC) MR images (1500/60; matrix, 128 x 128; four signals acquired) reveal enhanced contrast in healthy human brain. (a) IDQC-encode gradient was applied along direction of the magnetic field (B0). (b) IDQC-encode gradient was applied along "magic angle," where signals are minimized. (c) Conventional T2-weighted single-quantum coherence image. Images in a and b are displayed with same window setting but are different from that for c. (Image courtesy of Jianhui Zhong, PhD, University of Rochester, Rochester, NY.)
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Figure 9b: Transverse intermolecular double-quantum coherence (IDQC) MR images (1500/60; matrix, 128 x 128; four signals acquired) reveal enhanced contrast in healthy human brain. (a) IDQC-encode gradient was applied along direction of the magnetic field (B0). (b) IDQC-encode gradient was applied along "magic angle," where signals are minimized. (c) Conventional T2-weighted single-quantum coherence image. Images in a and b are displayed with same window setting but are different from that for c. (Image courtesy of Jianhui Zhong, PhD, University of Rochester, Rochester, NY.)
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Figure 9c: Transverse intermolecular double-quantum coherence (IDQC) MR images (1500/60; matrix, 128 x 128; four signals acquired) reveal enhanced contrast in healthy human brain. (a) IDQC-encode gradient was applied along direction of the magnetic field (B0). (b) IDQC-encode gradient was applied along "magic angle," where signals are minimized. (c) Conventional T2-weighted single-quantum coherence image. Images in a and b are displayed with same window setting but are different from that for c. (Image courtesy of Jianhui Zhong, PhD, University of Rochester, Rochester, NY.)
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MR imaging is a vibrantly evolving technology, and its influence on clinical medicine and patient care continues to expand. It can be argued that the increasing flexibility, wider applications, declining costs, more subtly nuanced pulse sequences, and growing magnetic field strength (ie, the increase from 1.5 to 3.0 T) of MR imaging, as well as the associated improvement in signal-to-noise ratio (65,66), are placing the technology in a position to displace many of the applications of CT, PET, and conventional angiography.
US Imaging
US produces images from high-frequency sound echoes that are created at boundaries between tissues with different elastic properties or densities (which, in turn, determine the speed of sound and the acoustic impedance for a medium). With its ability to do so in everyone from fetuses to the elderly and with no exposure to ionizing radiation, US is indeed a "womb-to-tomb" clinical imaging modality.
Over recent years, the evolution of diagnostic US has been quietly spectacular (67–69). There have been major strides in the design of miniaturized piezoelectric and capacitive transducers; contrast agents (which, at present, consist of suspensions in a fluid of erythrocyte-sized gas bubbles that may be coated with a tissue-targeting compound) (70,71); the exploitation of harmonic information, which is associated with higher frequency components created from nonlinear propagation of sound through tissues or nonlinear oscillations of contrast agent gas bubbles (72); the use of coded-excitation pulse trains and matched filters; image-processing software; techniques based on the absorption or speed of ultrasound radiation, rather than just on reflections; and small size and portability. These advances allow, with two-dimensional arrays of piezoelectric elements, the creation of real-time images that appear truly three-dimensional and that display high spatial, contrast, and temporal resolutions over a large field of view (73,74).
Clinical US systems normally operate in the 2–15-MHz range. Despite the technical difficulties, devices are now being designed with higher frequencies, with the shorter wavelengths allowing the imaging of correspondingly smaller objects. Micromachined US transducers, consisting at present of tens of separate piezoelectric or capacitive elements a few millimeters across—fine enough to fit within a narrow-gauge catheter and then into a coronary artery—are being built by way of microelectromechanical systems semiconductor manufacturing technologies (75). Vascular imaging is currently being performed in the 10–40-MHz range; at even higher frequencies (up to several 100 MHz), there are potential applications in ophthalmology, dermatology, and perhaps cellular-level imaging.
There are also early but promising results for true four-dimensional tomographic US imaging (Fig 10). US tomography may require hundreds of transducer elements covering a large-angle area of the body and perhaps firing and gathering echoes separately—although it is always possible that a much simpler solution to four-dimensional image acquisition may be found. Transmission US tomography has been demonstrated in relatively homogeneous tissues such as the female breast, where (as in refraction seismography, used by geophysicists in the search for oil and gas) variations in the speed and/or absorption of sound are measured and processed to create images. Reflection US tomography, with its multiplicity of omnidirectional echoes, continues to pose major technical problems that remain intractable at this time.
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Figure 10: Four-dimensional US images of breast in a 39-year-old woman with stage T2 cancer (arrow) in the left breast. Left: Cancer initially measured 24 x 12 x 19 mm. Patient was examined with combined three-dimensional US and digital mammography research system being developed at the University of Michigan (Ann Arbor, Mich) in partnership with GE Global Research. Right: Scans obtained after completion of four cycles of neoadjuvant chemotherapy. By viewing nearly whole breast image volumes, it is easier to localize any posttreatment remains of tumor and to judge treatment success, hopefully with automatic volume change measurements. Such four-dimensional imaging may revolutionize three-dimensional US screening for breast cancer, allowing precise section-by-section or pixel-by-pixel visual comparison. Ant = anterior, Lat = lateral, Med = medial, Post = posterior. (Image courtesy of Charles Meyer, PhD, and Paul Carson, PhD, University of Michigan, Ann Arbor, Mich)
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It is discontinuous changes in elastic properties or density at tissue interfaces or within tissues that are responsible for the creation of ultrasound echoes. The images generated, however, reveal little about tissue elasticity itself, which may nonetheless be of clinical interest. Elastography, or elastic imaging, reveals tissue elasticity parameters directly: US images generated before and after the mechanical or acoustic application of static or dynamic pressure or shear to effect small amounts of compression or strain are compared (76). With the related technique of MR elastography, the tissues of interest are stressed with small (10–100-µm amplitude) vibrations, and the induced distortions are measured with a phase-contrast pulse sequence (77).
A new area of physics, known as time-reversed acoustics, provides the means to detect a sound wave from a submarine, for example, and theoretically time-reverse the sound wave and send it back to its origin. This ability to locate the source may lead to novel ideas for both diagnostic and therapeutic applications of ultrasound (78).
Finally, there is active research on biologic effects (79), especially as US applications move to much higher frequencies. This is ongoing not only to ensure patient safety in diagnosis but also to support novel applications in therapy (80).
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DEVELOPING AND EXPERIMENTAL TECHNOLOGIES
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In this section, we will consider some of the newer technologies that are currently undergoing rapid development or are in an experimental stage and that are likely to be familiar to radiology clinicians and researchers and that, in the authors' opinion, demonstrate promise of success. A longer article would have mentioned a number of others, as well.
Gamma rays and x-rays reside at the top of the useable electromagnetic spectrum, and those fields are well plowed, as already indicated. Heading downward in energy, after meandering through the ultraviolet region (where currently there are few applications), one comes to visible and near-infrared light.
Optical and Near-Infrared Imaging
Diaphanography, also termed transillumination, is the medical counterpart to shining a bright flashlight through the hand to display the bones. As with x-rays, beams of visible and near-infrared light are differentially attenuated as they transit different types and thicknesses of tissues, and their shadows can be recorded as images on film or with an electronic camera. Transillumination of the breast, for example, can sometimes help distinguish benign from malignant masses—but this modality produces more false-positive and false-negative results than do conventional mammography and US, and it has not been accepted as a standard clinical tool.
The clinical utility of optical methods depends largely on the ability to extract image information about objects embedded in turbid media, so there have been serious efforts to reduce the level of scattered radiation reaching the image receptor. Improvements in time-resolved spectroscopy techniques and in the theory of photon transport in tissues have both contributed to recent advances in the field. In the former category, one family of methods employs gated detectors that are fast enough that, following an extremely short laser pulse, only unscattered photons (which travel directly from source to detector and arrive first) are accepted. This approach can display tumors and other abnormalities by recording light transmission as a laser is stepped through a region of interest—a much more refined version of diaphanography.
Laser optical tomography or, for one particular application, tomographic laser mammography, yields cross-sectional images of tissue obtained by projecting laser beams inward from many directions. These images are useful in studies of blood perfusion, tissue oxygenation, and neovascularity in the brain, breast, and extremities. In one interesting variation on this idea, photographic-acoustic tomography makes use of differences in the tendencies of tissues to absorb brief radiofrequency pulses of laser light and to heat up and thermally expand extremely rapidly, thereby producing ultrasonic waves that can be detected with sensitive piezoelectric receptors.
With a somewhat different approach, confocal scanning laser tomography can be used to noninvasively acquire three-dimensional images of the posterior segment of the eye, creating a quantitative description of the optic nerve head and the surrounding retinal surface (Fig 11). A laser beam is focused to some depth within the eye and scans a two-dimensional plane. Only light from that focal plane is allowed to reach the detector. A sequence of such two-dimensional optical planar views is acquired for increasing depths of the focal plane, and the result can be displayed as a three-dimensional topographic image of the optic nerve head and peripapillary retinal nerve fiber layer. The approach is useful not only for detecting abnormalities directly but also for tracking subtle changes over time that would be difficult to detect clinically. Authors of a recent study (81) suggest that, with expert interpretation, the power of the technique as an aid in the diagnosis and management of glaucoma is comparable to that of the accepted standard of stereofundus photography.
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Figure 11: Confocal scanning laser tomographic image of posterior segment of the eye shows topography of optic nerve head and of surrounding retinal surface in three dimensions. In less than 2 seconds, the 670-nm laser performed a sequence of 64 scans of the retina over a 15° x 15° field, creating 64 386 x 386-pixel planar images out of light reflected from different depths. The shape of the indentation edge, emphasized by the drawn green line, indicates nerve fiber layer defect at the rim of the optic nerve head. (Image courtesy of Heidelberg Engineering, Dossenheim, Germany.)
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Optical techniques such as optical coherence tomography and phase-resolved microscopy permit real-time micrometer-scale imaging and very sensitive display of cellular dynamics to depths of several millimeters in tissue. In optical coherence tomography, a thin beam of pulsed optical or infrared laser light passes into an interferometer, one arm or which is directed at tissue through an optical fiber, which is, perhaps, part of an endoscope (82). Light that scatters coherently in the top few millimeters of tissue contains information not only on the relative absorption of the radiation but also on changes in the light's phase. Scattered radiation returning from the tissue is caused to interfere with light from the reference arm, so that the phase shifts and the degree of absorption can be measured. When such interference data are obtained over a 1024 x 1024-pixel matrix imposed on the tissue surface, the system can form a 1-megapixel cross-sectional image in real time. Since the time of arrival of the scattered photons at the detector depends on the depth of scatter within the tissue, different levels in the tissue can be examined separately. Penetration is only several millimeters, but resolution can be 10 µm or better. Combination of the images with spectroscopic information can provide even more detail about the tissue for possible use for optical biopsy.
Diffuse optical tomography provides measurements of hemodynamics and neural activation at depths of several centimeters in tissue. Nonlinear microscopy employs nonlinear optical methods, such as multiphoton molecular excitation, optical harmonic generation, and depletion of stimulated emission. Nonlinear microscopy can be used to image subcellular morphology and trace molecular dynamics at subnanometer resolution at depths of up to a fraction of a millimeter in living tissue.
Tissue is nearly opaque throughout most of the infrared, visible, and ultraviolet parts of the electromagnetic spectrum. One way to deal with this challenge is to use a very intense laser beam, as in diaphanography. Another approach is to exploit a narrow window of transparency in the near-infrared region, with wavelengths in the 700–900-nm range. An important set of applications makes use of the fact that oxygenated and deoxygenated hemoglobin both absorb these wavelengths well but have spectra that differ enough to allow distinction between the two species. This distinction can be important in studies of blood flow and oxygen consumption.
Each year, 1 million women undergo core-needle breast biopsy in the United States. While considerably easier on the patient than excisional biopsy, the core-needle technique suffers from a false-negative rate of up to 7%, despite needle guidance with x-ray fluoroscopy or US. A more sophisticated type of probe currently under development may reduce the false-negative rate substantially (83,84). After the probe has been guided to the tumor site within the breast, it transmits near-infrared laser light into the breast tissue and then senses the light coming back from the tissue. The light returning from a tumor may differ from that returning from normal tissue because of irregularities in the degree of oxygenation, fluorescence characteristics, or other properties. In cases where no difference is detected, the tip of the probe can be repositioned in further search for disease. On finding tissue that yields a difference in returning light, the probe is used to acquire a core biopsy specimen. Unlike mammograms, the laser probe works well in dense tissue, such as in the breasts of young women, and may find applications in other organ systems, as well.
Terahertz Imaging
The terahertz portion of the electromagnetic spectrum lies between 300 and 100 µm in wavelength. So-called terahertz rays, or T rays, do not penetrate water or tissue well, so they would be of little use in the examination of deep-seated tissues. A substantial fraction of cancers lie in the epithelium, however, and while many of these are readily apparent to the trained eye, some that are small and flat can be overlooked or are simply not visible. Standard modalities are not adept at depicting or characterizing epithelial tumors—but terahertz imaging, because of the ability to recognize the spectral fingerprints of surface proteins that are markers for certain cancers, seems capable of demonstrating them at an early stage when they can still be treated effectively (Fig 12). Terahertz rays are nonionizing and offer imaging resolution of less than a millimeter; the equipment is safe and portable. There are major difficulties with terahertz sensing and imaging systems: They are inefficient and extremely costly to produce, but recently created photonic band-gap materials may soon radically change that situation (85–87).
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Figure 12a: Terahertz imaging. (a) In vivo terahertz image of volunteer's forearm acquired by using a handheld terahertz imaging system. The image shows a 15 x 15-mm region; a scar running from left to right can be seen in the top half of the image. The image was generated by plotting the electric field value reflected from beneath the skin surface. Dark circular regions are hair follicles of normal skin, which are not present at the scar. (b) Axial terahertz image (b scan) of edge of volunteer's hand. Gray scale indicates signal amplitude, which is plotted against optical delay (y-axis) and position across the scanned area (x-axis). Decrease in stratum corneum thickness across the x-axis, from the palm-side (30 mm on x-axis) to the backside (70 mm) of the hand, is evident. (Image courtesy of Vincent Wallace, PhD, TeraView, Cambridge, England)
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Figure 12b: Terahertz imaging. (a) In vivo terahertz image of volunteer's forearm acquired by using a handheld terahertz imaging system. The image shows a 15 x 15-mm region; a scar running from left to right can be seen in the top half of the image. The image was generated by plotting the electric field value reflected from beneath the skin surface. Dark circular regions are hair follicles of normal skin, which are not present at the scar. (b) Axial terahertz image (b scan) of edge of volunteer's hand. Gray scale indicates signal amplitude, which is plotted against optical delay (y-axis) and position across the scanned area (x-axis). Decrease in stratum corneum thickness across the x-axis, from the palm-side (30 mm on x-axis) to the backside (70 mm) of the hand, is evident. (Image courtesy of Vincent Wallace, PhD, TeraView, Cambridge, England)
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Microwave Imaging
Microwave imaging is particularly sensitive to differences in absorption between fat and other soft tissues. The wavelength is too long to allow the degree of spatial resolution normally required for diagnostic imaging, but microwave devices may find other applications (85,88). For example, promising preliminary clinical studies have been reported in which a small device called a tissue resonance interferometer was used. This device generates low levels of 400–1350-MHz radiation (less than that of a cell phone), and the returning signal is altered by tissue irregularities, including tumors.
Electron-spin resonance imaging, also known as electron paramagnetic resonance imaging, is analogous to MR but involves spin transitions of unpaired electrons rather than of (unpaired) protons (60,89). Since an electron is three orders of magnitude lighter than a proton, it has a considerably greater magnetic moment and a correspondingly higher Larmor frequency. At 1 T, an electron's magnetic resonance occurs at 28 GHz, rather than at 42 MHz for a proton (1 GHz = 1000 MHz). Microwave energy tends to be strongly absorbed in the process of excitation of rotational states in water—indeed, that is how microwave heating works—and at gigahertz frequencies, microwaves are absorbed in a millimeter of skin. At considerably lower resonance frequencies and field strengths (eg, 0.01–0.04 T), however, the penetration permits imaging of free radicals in vivo in small animals. A variation on the theme, known as proton-electron double-resonance imaging, or PEDRI, involves the simultaneous performance of electron-spin resonance and nuclear MR imaging (Fig 13) (90).
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Figure 13a: Proton electron double resonance imaging (PEDRI; also known as Overhauser imaging). (a) Time course of PEDRI study of myocardial uptake of free radical probe TEMPONE (4-oxo-2,2,6,6-tetramethylpiperidine-1-oxyl) by isolated perfused rat heart. TEMPONE was infused through a side arm proximal to the perfusion cannula at final concentration of about 3 mol/L. Two-dimensional PEDRI sections were then sequentially acquired every 30 seconds, with each scan taking 27 seconds. At low field strength of 0.02 T (201 G), the electron-spin resonance and nuclear MR frequencies are 567 MHz and 856 KHz, respectively, and an Overhauser enhancement of –13 was achieved. (b) Three-dimensional gradient-echo PEDRI images of isolated beating rat heart infused with 3 mol/L TEMPONE. Top left image shows complete three-dimensional surface-rendered image; the other images are cutaways to show internal structure. The image took 4 minutes 30 seconds to acquire at 0.02 T. (Images courtesy of Haihong Li and Jay Zweier, MD, Davis Heart and Lung Research Institute, Ohio State University, Columbus, Ohio.)
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Figure 13b: Proton electron double resonance imaging (PEDRI; also known as Overhauser imaging). (a) Time course of PEDRI study of myocardial uptake of free radical probe TEMPONE (4-oxo-2,2,6,6-tetramethylpiperidine-1-oxyl) by isolated perfused rat heart. TEMPONE was infused through a side arm proximal to the perfusion cannula at final concentration of about 3 mol/L. Two-dimensional PEDRI sections were then sequentially acquired every 30 seconds, with each scan taking 27 seconds. At low field strength of 0.02 T (201 G), the electron-spin resonance and nuclear MR frequencies are 567 MHz and 856 KHz, respectively, and an Overhauser enhancement of –13 was achieved. (b) Three-dimensional gradient-echo PEDRI images of isolated beating rat heart infused with 3 mol/L TEMPONE. Top left image shows complete three-dimensional surface-rendered image; the other images are cutaways to show internal structure. The image took 4 minutes 30 seconds to acquire at 0.02 T. (Images courtesy of Haihong Li and Jay Zweier, MD, Davis Heart and Lung Research Institute, Ohio State University, Columbus, Ohio.)
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Thermography
In contrast to the modalities discussed so far, thermography, electrocardiography, electroencephalography, magnetoencephalography, and magnetocardiography all involve extraction of information from the electromagnetic radiation that the body itself produces and emits naturally.
The warmer something is, the more heat it emits, at a rate that is approximately proportional to the fourth power of the absolute temperature (ie, approximately T4). Like some night-vision techniques, emission thermography has long been used to sense infrared radiation (which has a lower frequency than visible light) emitted through the skin as a result of heat brought to the surface by blood flowing from deeper regions. Irregularities that affect blood flow in the outermost few millimeters of the skin or that influence its temperature directly (eg, angiogenesis and neovascularity that accompany breast disease) can be detected by means of an infrared camera.
While there is not a great deal of evidence supporting the relative value of thermography in helping search for breast tumors, it has proved to be effective in monitoring of inflammatory conditions such as rheumatic disease, injured muscles (Fig 14), diminished enervation of muscles, burns, and frostbite (91).
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ABSTRACT
INTRODUCTION
