© 2004 American Thoracic Society DOI: 10.1165/rcmb.2003-0213TR
Recent Advances in Imaging the Lungs of Intact Small AnimalsDepartments of Medicine and Radiology, Washington University School of Medicine, and Department of Chemistry, Washington University in St. Louis, St. Louis, Missouri Address correspondence to: Daniel P. Schuster, M.D., University Box 8225, Washington University School of Medicine, 660 S. Euclid, St. Louis, MO 63110. E-mail: daniel.schuster{at}wustl.edu
A new generation of imaging devices now make it possible to generate both structural and functional images for the study of lung biology in small animals, including common laboratory mouse and rat models. "Micro" X-ray computed tomography and positron emission tomography scanners, highly sensitive cooled charge coupled device cameras for bioluminescence and fluorescence imaging, high magnetic field magnetic resonance imaging scanners, and recent advances in ultrasound system technology can be used to study such diverse processes as ventilation, perfusion, pulmonary hypertension, lung inflammation, and gene transfer, among others. Images from more than one modality can also be fused, allowing structure–function and function–function relationships to be studied on a regional basis. These new instruments, part of an emerging suite of techniques collectively known as "molecular imaging," provide an enormous potential for elucidating lung biology in intact animal models and systems.
Abbreviations: apparent diffusion constant, ADC • charge coupled device, CCD • computed tomography, CT • functional residual capacity, FRC • green fluorescent protein, GFP • magnetic resonance imaging, MRI • near-infrared, NIR • positron emission tomography, PET • radio-frequency, rf
A group of new techniques, known collectively as "molecular imaging," now offer scientists an unprecedented opportunity to identify, follow, and quantify biologic processes at the cellular and subcellular level in intact organisms. For instance, it is now possible to evaluate, with imaging, the distribution, magnitude, and timing of gene expression in genetically altered animals (1, 2). Even though nonimaging methods are also available to monitor gene expression, such methods—usually based on tissue sampling—are invasive and unattractive as routine procedures for clinical investigations. In contrast, molecular imaging can provide a seamless translation from studies in animals to later studies in humans. Although most studies that employ molecular imaging have so far focused on non–organ-specific applications (such as cancer detection or treatment monitoring and gene therapeutics), the ability to image fundamental processes (such as gene expression, inflammation, cell trafficking, apoptosis), provides ample reason to employ these methods in studies of lung biology. This rapidly developing, multidisciplinary field capitalizes on recent advances in the techniques of molecular and cell biology, on new highly specific probes that serve as sources for the imaging signal, and on dramatic improvements in imaging instrumentation (especially for small animals). In this brief review, we highlight and illustrate some of these new technologies. In-depth, general reviews of molecular imaging are available elsewhere (1, 3–8). Here, we focus on the types of lung studies that are being conducted (or could be), and on what the current limitations are for imaging lung structure or function (and the prospects for near-term improvements).
The value of X-ray–computed tomography (CT) is familiar to everyone involved in the clinical practice of pulmonary medicine. An enormous amount of useful anatomo-spatial information is generated from measured differences in tissue absorbance of X-radiation as it passes through the body from an external source. These differences in tissue absorbance, especially for soft tissues such as the lungs, are largely the result of variations in tissue density. Thus, in essence, CT images of the lungs are two-dimensional maps of tissue density. The ability to quantify lung density can be exploited to reveal insights into lung function as well as lung anatomy. For instance, the functional residual capacity (FRC) is an important determinant of gas exchange. Given measurements of lung density over the entire thorax, it is possible to calculate the volume of gas within the lungs in a straightforward manner. With recent improvements in the spatial resolution capacities of CT scanners, these principles have been used to extract FRC data from CT scans in animals as small as mice, even using clinical scanners designed primarily for human use (9–11). Interestingly, the FRC measurements in these studies were obtained without respiratory gating or breath-holding. Mitzner and coworkers have speculated that the accuracy of the FRC estimates was the result of off-setting errors, in which the respiratory motion artifact due to breathing was on the same scale as the spatial resolution of the clinical CT scanner (9). Generally, the density of an organ is relatively fixed, but within the thorax, lung density changes significantly, even dramatically, as a function of the respiratory cycle. Thus, when lung density measurements are correlated with coincident measurements of inflation pressure, the compliance of the lung and/or thoracic cage can be calculated to study the mechanical properties of these structures. The recent development of high-resolution CT imaging has made it possible to manufacture scanners specifically designed for imaging small animals ("microCT"). Relevant issues, such as X-ray source, focal spot size, detector element size, system geometry, and X-ray flux, are important in determining image spatial and contrast resolution, and are discussed in detail elsewhere (12). State-of-the-art systems can now produce images with a spatial resolution of 50–100 µ (2). An example of the detail that can now be achieved with one of these systems is shown in Figure 1. At such levels of resolution, detailed measurements of airway caliber, even in rodents, should be possible.
One weakness of CT (for instance, compared with magnetic resonance imaging) is its relatively poor soft-tissue contrast. Thus, for many clinical applications, contrast-enhancing media are injected to help distinguish structures such as blood containing vessels or cardiac cavities from surrounding soft tissues. Limited reports suggest that some degree of contrast enhancement may also be possible during microCT imaging (12). In the lungs, the relatively radiodense gas, xenon, has also been used as a contrast-enhancing agent to visualize airways. When intra-airway density measurements are made over time, it is possible to derive estimates of regional lung ventilation (10). To date, no such reports with xenon gas have been reported for microCT studies. Another issue of some importance is the radiation dose delivered to the animal during CT imaging. This may approach 5% of the LD50 in mice, potentially limiting the number of repeat studies that could be performed over time (12). Despite the possibility of using CT for functional imaging, its principle use is likely to still emphasize anatomic over physiologic applications. For small animal imaging, the ability to overlay anatomic displays obtained by CT with functional molecular imaging obtained by PET or MRI will be especially important to localize physiologic events identified by these latter two modalities (see below).
Proton (1H)-Magnetic Resonance Imaging Magnetic resonance imaging (MRI) is another powerful and versatile imaging modality for noninvasive, in-vivo characterization of both lung structure and function. To date, most MRI experiments involve measuring the protons of water, which are ubiquitous and present in high concentrations in animal tissues and organs. The fundamental principle underlying 1H-MRI is that when an individual, lying within a magnetic field, is subjected to a radio-frequency (rf) pulse applied at the correct ("resonant") frequency, the protons within the body absorb energy and generate a detectable signal. The signal strength depends upon the number (concentration) of protons; its frequency depends on the identity of the nuclide (e.g., 1H, 3He) and the strength of the local magnetic field. By applying magnetic field gradients along well-defined directions in space, the frequency of the proton signal is made dependent upon spatial position. Collection and processing of this signal produces a map of MR signal intensity versus position, i.e., an image. Following excitation by rf pulses, protons (and all other MR-active nuclides) return to equilibrium ("relax"). This relaxation process can be described by two fundamental rate constants: longitudinal relaxation (T1) and transverse (T2) relaxation. A detailed treatment of these two parameters is beyond the scope of this review. However, the important points for MRI are: (i) different tissues can have different T1 and T2 values; (ii) the intensity of signals in MRI images can be made sensitive to these relaxation parameters. As a result, T1 and T2 become potential sources of contrast between different organs within the body and between healthy and diseased tissue. MR imaging involves the sequential acquisition of signals in a series of experiments designed to map three-dimensional space. MRI experiments can be categorized into two basic types: gradient-echo experiments, which use a single rf pulse, and spin-echo experiments, which use two rf pulses. Gradient-echo experiments can generally be performed more quickly; spin-echo studies generally produce higher quality images. All MR imaging experiments are described by two fundamental timing parameters, TR and TE. TE is defined as the time between the excitation rf pulse and the start of data acquisition, whereas TR is defined as the time between sequential acquisitions. T1- and T2-weighted imaging refers to experiments designed to generate image contrast by taking advantage of differences in these relaxation parameters between tissues. T2-weighting is achieved through adjustment of TE, T1-weighting by changes in TR. Imaging the lung parenchyma, especially in small animals, presents some unique challenges for MRI (13), including: (i) relatively low tissue density (and therefore low water content within the lungs), severely limiting signal-to-noise; (ii) variations in magnetic "susceptibility," associated with the many air–tissue interfaces of the alveoli and bronchioles, creating local magnetic field inhomogeneities (field gradients) that can lead to shortening of some relaxation times; and (iii) respiratory and cardiac motion, leading to significant image blurring. In response to these challenges, researchers generally use very fast gradient echo imaging sequences with short TE times or spin-echo sequences with respiratory synchronization (14, 15). In addition, static magnetic field strength, gradient strength, and rf coil performance can all affect sensitivity and spatial resolution in MRI experiments. For small animal imaging studies, 4.7 T MR scanners are commonly used (compared with the 1.5 T magnetic field typical of clinical scanners). High gradient strengths (up to 60 G/cm, compared with 2 G/cm on most clinical scanners) are also typical. Experiments in small animals are performed with respiratory gating (15, 16), because breath-holding is often not an option. Data collection varies from minutes to one or more hours/animal, depending upon the nature of the experiments being performed. Many pulmonary studies in small animals have focused on edema (17, 18) or other disease states, including lungs affected with tumors or fibrotic lesions, in which the density of tissue increases and the corresponding susceptibility broadening due to air–tissue interfaces is decreased. In general, a good correlation has been reported between MRI-determined estimates of lung water and the gold-standard gravimetric method (13, 19). MRI has been used to assess edema and inflammation in the lungs following allergen or endotoxin challenges (20, 21), after exposure to an 85% O2 atmosphere, and after treatment with paraquat (15). Figure 2 shows respiratory-gated spin-echo 1H-MR images of a healthy mouse and one whose lungs are filled with fibrotic lesions after administration of intratracheal bleomycin.
Hyperpolarized Gas Imaging As with xenon-enhanced CT imaging, airway imaging with MRI can also be enhanced by the administration of an inhaled gas. MRI of normal gases suffers from very poor signal-to-noise due to the extremely low density of gas. However, with hyperpolarized noble gases (22), such as 3He and 129Xe (23), magnetization can be 105 that at thermal equilibrium. These extremely high nuclear spin polarizations more than compensate for the 103 lower density of the gaseous-state relative to that of liquid water, yielding images that are, in many cases, of higher sensitivity than observed in 1H-MR images of tissue water. In its most straightforward application, hyperpolarized gas can be used as a gaseous inhaled contrast agent. Regions of the lung that are well ventilated will appear bright in images following inhalation of hyperpolarized 3He; regions of impaired ventilation appear darker (23–26). Measurement of signal intensity in a single, multi-slice set of rapidly collected images produces a ventilation map. T1 relaxation of 3He in lungs is due almost entirely to interaction with oxygen in the gas phase. However, given the relatively small variations in PO2 in lungs(27), the effects of minor variations in T1 on image intensity can be minimized by collecting data on a time scale that is short compared with typical 3He T1s (20 s). In humans, these methods have been used to study a wide variety of disease states, including chronic obstructive pulmonary diseases (COPD) such as emphysema (24, 28), asthma (29), cystic fibrosis (30), and bronchiectasis (31). Similar experiments can be performed in vivo in small animals. However, imaging of small animals requires smaller voxels, so that a single breath-hold of hyperpolarized 3He (as typically performed in clinical studies) does not yield adequate signal-to-noise. Instead, small volumes of hyperpolarized gas are repetitively delivered using a ventilator and data are collected over periods of many minutes. Figure 3 shows one slice from a multi-slice, hyperpolarized 3He ventilation image of a mouse lung collected in the laboratory of one of the authors (J.G.). Several recent studies have focused on generating high-resolution lung ventilation maps in small animals using hyperpolarized 3He gas (31–34), whereas Johnson and colleagues recently described a combined 3He/1H imaging study of rat lung (35).
Dynamic imaging experiments using hyperpolarized gas, in which a series of time-resolved intensities are measured during the inspiration/expiration cycle, can provide information about functional lung ventilation and ventilation distribution that is absent from static imaging experiments (36). Dynamic imaging is achieved by selecting experimental conditions to observe either recently inhaled, fresh gas only, or all of the gas within the lung. Figure 4 shows a set of serial rat lung images acquired during inhalation of hyperpolarized 3He. It is also possible to obtain a simultaneous regional assessment of both lung ventilation and perfusion, by monitoring 3He image intensity as a function of time following an intravenous bolus injection of superparamagnetic iron oxide nanoparticles (37, 38).
Diffusion MRI measurement of hyperpolarized gas diffusion is another important source of information about lung morphology and disease. The diffusion of gas is restricted by the microstructure of the lung—its small airways and alveolar walls—making diffusion sensitive to microstructural features that are at an alveolar distance scale, well below the spatial-resolution limits of conventional MR imaging. These properties can be quantified as a gas apparent diffusion constant (ADC), complementing information about the uniformity of gas transport derived from ventilation images. In healthy lungs, gas diffusion, as measured by ADC, is homogeneous, reflecting the uniformity of tissue within the lungs. A typical ADC value in the alveolar spaces in healthy rodent lung is 0.16 cm2/s, more than an order of magnitude smaller than the ADC in trachea (2.4 cm2/s), where diffusion is relatively unrestricted (15). In an elastase-induced model of emphysema in rats, Chen and associates recently demonstrated that the breakdown of elastic components leads to an overall enlargement of alveolar volume, lessening restrictions to diffusion and increasing diffusivity as measured by ADC (39). The local cylindrical geometry of the airways not only reduces the average diffusion rate, but also produces local orientation-dependent effects, with diffusion along (parallel to) the long axis of an airway being significantly faster than diffusion across (radial or perpendicular to) the long axis. Yablonskiy and coworkers recently reported that diffusion in acinar airways is anisotropic, and demonstrated a significant difference in anisotropy between healthy and emphysematous lungs (40). In healthy lungs, an anisotropy-derived estimate of the mean radius for acinar airways is in excellent agreement with results of histology on excised lungs. These methods can now be extended to longitudinal studies characterizing changes in mean small-airway radius in animal models of lung disease. Both X-ray CT and MRI can be used to measure morphologic properties such as airway caliber and functional assessments of ventilation and perfusion. Although X-ray CT imaging of the lung parenchyma is still often superior to that obtained by MRI, MRI has the advantage of being able to generate images without potentially damaging ionizing radiation. This may be especially important for small animal imaging studies conducted over multiple imaging sessions, as the radiation dose of X-ray CT in small animals can be considerable (see above). On the other hand, there may be many instances in which the unique advantages of both modalities can be combined, especially with new display software that allows images to be overlaid onto one another.
Positron emission tomography (PET) is a powerful, quantitative, nuclear medicine imaging technique, which can be used to study many interesting and important problems in lung physiology and biochemistry, such as peptide metabolism, blood flow, ventilation, and water content (41–43). More recently, PET imaging methods have been reported which can be used to evaluate lung inflammation and pulmonary transgene expression (44, 45). After compounds are labeled with positron-emitting isotopes, they are administered intravenously or inhalationally, and the tissue activity concentration of the isotope is determined with an imaging device (the "PET scanner"). Multiple two-dimensional images are then reconstructed from the activity data and interpreted mathematically to represent the process of interest. PET derives its power from several factors: (i) the labeled compounds themselves are often biologically important (e.g., water); (ii) the isotope half-life is often sufficiently short that studies may be repeated if desirable; (iii) the regional isotope concentration can be determined quantitatively, accurately, and noninvasively; and (iv) the activity distribution can be located with increasingly improved spatial resolution. Given this latter feature in particular, the activity data can be presented regionally, in an image format, so that measurements may be correlated with other regionally specific measurements over time. The isotopes used in PET all decay by a process of positron emission. The emitted positron is quickly annihilated by its interaction with an ambient electron. Two photons are emitted after each annihilation event, traveling nearly 180° apart. These photons may be counted with detectors placed on opposite sides of the body. Because the annihilation event that produced the photons can be assumed to have occurred somewhere within the tissue volume subtended by the two detectors, the radiation source can be partially located in space. As additional detector pairs are added to the system, intersecting lines further establish spatial location. This type of collimation by electronic coincidence detection is far more efficient than that provided by the lead shielding used in conventional single photon emission detection systems because less radiation is ignored. The richness of using PET to study lung physiology and metabolism is based on the specificity of radiolabeled probes which can be harmlessly injected or inhaled. With tracers labeled with oxygen-15 (t1/2 = 2.1 min), carbon-11 (t1/2 = 20 min), or nitrogen-13 (t1/2 = 11 min), repetitive studies are readily performed. Among the simplest tracers used in pulmonary PET studies are [15O]H2O, [15O]CO, and [13N]N2. These natural compounds are used for regional perfusion, water content, and blood volume studies (41). More complex radiopharmaceuticals, such as labeled amino acids and amines, using carbon-11 or fluorine-18, can be useful in enzyme and receptor studies. These various features create obvious incentives for carrying out PET studies in small animals typically used for laboratory study. Until recently, limitations in spatial resolution made such studies impossible. For instance, most PET scanners in clinical use have an image spatial resolution of 10–15 mm, but the mouse thorax, in vivo, is only 20–25 mm wide. Other issues, such as scanner sensitivity (i.e., the fraction of radioactive events actually detected by the device) and the amount of radioactivity that can be injected without causing physiologic disturbances in the system under study (a function of tracer specific activity) also affect the ability to perform imaging studies in small laboratory animals. Nevertheless, recent advances in the scintillation materials used to make the radiation detectors used in PET devices, and in the ability to transfer the scintillation light from the detectors to photomultiplier tubes (4), now make it possible to manufacture devices with remarkable improvements in spatial resolution. An example of an image generated with a so-called "microPET" scanner is shown in Figure 5.
One recent, and particularly exciting, form of molecular imaging seeks to follow the expression of specific genes via the noninvasive visualization of in vivo gene "reporters" (46–49). Noninvasive real-time analysis of gene expression is possible using reporter genes with optical signatures (e.g., green fluorescent protein [gfp], firefly luciferase [50], and bacterial luciferase) (see following section). However, spatial resolution, quantitation, and the imaging of deep tissues is poor or problematic. Gene expression imaging with magnetic resonance (MR) is also possible, but this approach is not suitable for lung studies because the proton density of the lung parenchyma is too low to generate a suitable MR image(51). Radiotracer imaging, however, with methods such as PET, has a high degree of sensitivity (level of detection approaches 10-11 M of tracer), and isotropism (i.e., ability to detect expression accurately regardless of tissue depth, unlike light-based techniques, which are largely limited to detection at the body surface). Thus, they are ideally suited to detect gene expression in deep organs such as the lungs. Recently, we showed that PET imaging of transgene expression is possible in rodent lungs (Figure 6).
Often some physiologic model, expressed in mathematical terms, is required to interpret the activity measurements collected during a PET study, especially when serially consecutive images are obtained over a period of time to generate tissue time–activity relationships. These models must incorporate such factors as tracer delivery to tissue, blood concentration, tissue uptake, metabolism, the recirculation of both metabolized and unmetabolized tracer, and the heterogeneity of tissues within the resolution volume (41, 52). In applying a mathematical model to a series of PET images, the image is converted from a display of regional activity concentration into a map representing the regional distribution of the physiologic process being studied (e.g., blood flow, or a rate constant of tracer transfer from one tissue compartment to another). It is this transformed "parametric" image that is then analyzed and interpreted biologically. Needless to say, the physiologic data finally analyzed from such a PET study will only be as accurate as the model used to calculate them. To the extent such models can be used successfully, PET becomes not only another form of tomographic imaging, but a powerful analytic tool for studying in vivo biology. Such quantitation, however, depends on being able to measure a so-called "input function", i.e., the time–tracer activity relationship in blood perfusing the lungs. Because spatial resolution is still too poor to allow sampling of the blood pool within the hearts of rodents by PET imaging, such quantitation will require direct blood sampling, an obvious technical challenge.
PET is only one radionuclide-based method currently being evaluated as a platform for gene expression imaging. Others include planar
Whole body imaging of light generated deep within the body from biochemical reactions and biological processes is now feasible in laboratory animals and potentially in humans, representing an interesting new avenue for molecular imaging. Optical properties of tissues limit the ability to detect light from weak light sources within the body. Fundamentally, constraints for optical imaging arise from the low levels of light emitted by internal sources and the high absorption and scatter of light traversing the body (1). To enable optical imaging in vivo, new sensitive instruments and imaging probes have recently been developed. Three general strategies have evolved for optical molecular imaging in vivo: use of endogenous fluorochromes; use of reporter genes that generate internal light from specific biochemical reactions (bioluminescence and fluorescent proteins); and use of injected optical contrast agents incorporating visible light fluorophores, near-infrared fluorophores, or activatable fluorophores (1, 3, 55). Optical imaging is relatively low in cost, highly versatile, and enables multichannel imaging through use of multiple probes with differing spectral characteristics (56). Here, we focus on new instrumentation that has enabled whole animal imaging, including lung imaging, through the use of bioluminescence and fluorescent proteins and other agents. Bioluminescence specifically refers to the enzymatic generation of visible light by living organisms (57–59). Bioluminescence has particular appeal as an approach for optical imaging in vivo because mammalian tissues do not emit significant levels of intrinsic bioluminescence, enabling images to be generated with remarkably high signal-to-noise. However, bioluminescence imaging requires that the gene encoding the bioluminescent reporter protein be cloned into expression cassettes in cells or tissues of interest. Although the most commonly used bioluminescent reporter for research purposes has been luciferase from the North American firefly (Photinus pyralis; FLuc), useful luciferases have also been cloned from jellyfish (Aequorea), sea pansy (Renilla; RLuc), corals (Tenilla), and several bacterial species (Vibrio fischeri, Vibrio harveyi) (60). FLuc catalyzes transformation of the substrate D-luciferin into oxyluciferin in an ATP-dependent process, leading to the emission of photons, whereas RLuc catalyzes oxidation of the substrate coelenterazine into coelenteramide in an ATP-independent reaction. Moreover, luciferases and their cognate substrates appear to be nontoxic to mammalian cells (61). The sensitivity of detecting these internal light sources is dependent upon many parameters, including the level of luciferase expression, the depth of labeled cells within the body (i.e., the distance that the photons must travel through tissue) and the sensitivity of the detection system. Key advances in detector technology for imaging low levels of light now enable optical imaging in living animals with charge coupled device (CCD) cameras. By converting light photons that strike silicon wafers into electrons, CCD cameras spatially encode the intensity of incident photons into electrical charge patterns to generate an image (62). For bioluminescence imaging, the thermal noise of the systems is reduced by super-cooling the CCD camera, and the camera is mounted in a light-tight box (thereby increasing the sensitivity of the camera to low intensity light typical of a bioluminescence experiment). For example, as few as 100 transduced cells can be detected after injection into the peritoneal cavities of severe combined immunodeficient (SCID) mice; less than 10,000 cells can be visualized in the lungs after intravenous injection of labeled cells (63). In syngeneic animal models of leukemia and lymphoma, tumor cells expressing FLuc can be detected with high sensitivity in internal organs like lung, liver, spleen, lymph nodes, and even within the bone marrow of BALB/c mice (63). Using this approach, the spatiotemporal trafficking patterns of lymphocytes within the body can be elucidated. Figure 7 shows an example of bioluminescence imaging of FLuc transduced leukocytes in the lungs of a mouse tumor model.
Bioluminescence is relatively simple to execute relative to PET or similar radionuclide-based imaging technologies because the substrates (D-luciferin for FLuc and coelenterazine for RLuc) are commercially available and readily prepared for injection into whole animals. High throughput screening is feasible because of short image acquisition times, capacity for repetitive imaging after short time intervals, and a platform for simultaneous acquisition of multiple animals at the same time. However, a key disadvantage of cooled CCD cameras and bioluminescence imaging is the limited and wavelength-dependent transmission of light through animal tissues. As a rule of thumb, there is an approximate 10-fold loss of photon intensity for each centimeter of tissue depth (64). Also, the images are surface-weighted, meaning that light sources closer to the surface of the animal appear brighter compared with deeper sources due to tissue attenuation properties (1). Another disadvantage is the current limitation of planar display instead of the tomographic or three-dimensional displays typically seen with X-ray CT or PET. Thus, bioluminescent images lack depth information. However, technological advances with rotating mirrors may enable tomographic bioluminescence imaging in the near future.
Luciferase reporters have been used in a variety of animal imaging applications (65). Bioluminescence enables monitoring throughout the course of disease, allowing localization and serial quantitation before killing the experimental animal (63, 66). Bioluminescence imaging can be used for real-time studies of cell trafficking (67), of various genetic regulatory elements in transgenic mice (68), and of in vivo gene transfer (58, 69, 70). Bioluminescence methods have been used to study the ubiquitin-proteasome pathway, the central mediator of regulated proteolysis in cells (71). Bioluminescence also has been used to study the spatiotemporal patterns of bacterial and viral infections and their treatment using genetically engineered bioluminescent pathogens and whole body imaging (72–74). In the lungs, bioluminescence has been used to study both nuclear factor- In contrast to bioluminescence imaging, during fluorescence imaging, an excitation light illuminates the subject and a CCD camera, usually less sensitive than the cooled CCD cameras used for bioluminescence, detects the emitted fluorescence. Cells can be labeled with fluorescently labeled antibodies or express green fluorescence protein (GFP) or color-shifted/enhanced variants (yellow, cyan, red, EGFP, etc.). Like bioluminescence, the procedure generally involves transduction of the gfp gene into cells or tissues. Exciting the protein with appropriate blue light and detecting the ensuing fluorescence with a CCD camera enables whole body imaging of GFP expression in living animals (77). Nonetheless, absorption and fluorescence of GFP in the visible region, coupled with increased light scattering, may hinder the use of GFP for high-resolution imaging of deep tissues in vivo (63). Recent advances in fluorescent proteins, however, include the engineering or discovery of variants with enhanced brightness, improved pH resistance, the ability to undergo photochemical color conversion, and red fluorescent emissions that likely will lead to new applications in animal models (78). Advantages of fluorescence imaging include the capacity for both living and fixed cells and tissues to be visualized, the lack of a substrate to generate the fluorescence phenomena, and low cost of the animal imaging instruments (78). However, like bioluminescence, the images are surface-weighted. Another disadvantage of fluorescence imaging is the higher background signal levels due to autofluorescence of tissues, especially in the blue and ultraviolet wavelengths compared with bioluminescence. Nonetheless, a variety of metastatic tumor and immune cell trafficking models have been successfully imaged in vivo with fluorescence imaging of GFP-tagged cells (77, 78). As noted earlier, optical imaging can also be performed after administration of exogenous optical contrast agents, synthesized with fluorophores having emission characteristics in the visible and near-infrared (NIR) spectrum (700–900 nm). There is significant interest in applications in the NIR spectrum due to the high tissue penetration of light in this range as well as minimized autofluorescence. Many fluorophores are available for coupling to biologically targeted molecules such as antibodies and peptides (79, 80). In general, fluorescein, indocyanine green (ICG), and porphyrin derivatives are used as optical contrast agents for in vivo imaging, localizing to pathologic tissues by a variety of targeting mechanisms (81, 82). For example, a carbocyanine-somatostatin receptor-avid peptide conjugate was selectively retained in tissues overexpressing somatostatin receptor type-2, a receptor subtype overexpressed in small cell lung cancer. Another approach with exogenous agents involves use of activated optical probes or optical beacons (83, 84). These reagents typically are graft copolymers (methoxy-polyethylene-glycol-derivatized poly-L-lysine) appended with cleavable peptides conjugated between the polymer and NIR fluorescent (NIRF) fluorophores such as tricarbocyanine and indocyanine green dyes. The spectral properties of these dyes produce self-quenching when the fluorophores are co-assembled on the polymer. However, upon proteolysis of the intervening target peptide sequence, the fluorophores are released into the surrounding environment enabling the generation of florescence signals. This approach has been used, for example, to study tumor matrix metalloproteinase and cathepsin activities in vivo (3, 83, 84). As with bioluminescence, most fluorescence imaging is planar with the fluorescence image projected over a gray-scale photograph of the object. However, new approaches to fluorescence-mediated tomography are under investigation (85).
Clinical echocardiography has been widely used and validated as a reliable and accurate tool for non-invasive estimation of pulmonary artery pressures and for evaluation of the structural and functional effects of pulmonary hypertension on the right heart. Recent advances in ultrasound system technology now allow acquisition of cardiac images with high spacial and temporal resolution necessary to study murine models of pulmonary hypertension. However, the small size and rapid heart rate of rodent hearts still make some of these measurements technically challenging and experimentally impractical, especially in mice. A modification to the Bernoulli equation (PRV-RA = 4V2) (where P = pressure in mm Hg, RV = right ventricular, RA = right atrial, and V = Doppler flow velocity) is the primary basis for estimating pulmonary artery pressure with echocardiography (actually, the peak systolic right ventricular – right atrial pressure gradient) by measuring the peak velocity of the tricuspid regurgitation jet that typically accompanies significant increases in pulmonary artery pressure. This approach was adopted by Jones and colleagues (86) in a rat model of pulmonary hypertension. However, they found that quantitatively useful tricuspid regurgitation was only detected in animals with severe pulmonary hypertension (tricuspid regurgitation velocities of 3 m/s or higher). Pulmonary hypertension also alters the dynamics of pulmonic valve motion and blood flow velocity profile in a characteristic way: the period from the Q wave of the electrocardiogram to the onset of pulmonic valve opening (pre-ejection period) is prolonged, the time from onset to maximum velocity of pulmonic flow (acceleration time) is shortened, the ejection period is decreased, and flow can cease at mid-systole (depending on the severity of pulmonary hypertension). M-mode recordings commonly used in human studies to evaluate pulmonic valve motion are technically difficult in rats and impossible to obtain in mice because the thickness of the valve leaflets is beyond the spatial resolution of current ultrasound systems. Doppler recordings from the parasternal short axis view of pulmonary flow velocities, however, are easily obtained in both rats and mice. Structural changes, such as right ventricular hypertrophy, can also be easily evaluated as a common response to increased pulmonary pressure. Depending on duration and severity (as well as concurrent valvular abnormalities such as tricuspid regurgitation), pulmonary hypertension may be associated with right ventricular chamber dilatation (Figure 8) and impaired systolic function. High-resolution echocardiography can also accurately detect small changes in right ventricular wall thickness (Figure 9) (86).
The small size of the mouse heart (  10 times smaller than the rat heart) represents a significant challenge for imaging right heart structures by echocardiography. Right ventricular wall thickness and chamber dimensions are at the limits of spatial resolution of currently available ultrasound systems. Although echocardiography of the mouse left ventricle has been widely used and validated, there are limited data on imaging of the right heart. MRI, with its higher spatial resolution and inherent three-dimensional analysis capabilities, may offer a more accurate alternative to echocardiography for the noninvasive evaluation of right ventricular structure and function in mice (87). These advantages, however, have to be balanced against the expense and relative lack of availability of MRI versus echocardiography.
From the foregoing, it is obvious that this new generation of devices will allow pulmonary scientists to study in vivo lung biology at an unprecedented cellular and molecular level. A particularly intense area of research focuses on fusion imaging. Because each of the modern molecular imaging technologies like PET, MRI, and optical imaging has specific strengths and weaknesses, it makes sense to identify ways to benefit from each. Fusion imaging will allow detailed structure–function imaging (e.g., with X-ray CT and PET). The use of multifunctional reporter genes that link two or more modalities is another approach that should be highly informative. Fusion genes that encode bioluminescent and fluorescent reporter proteins effectively couple the powerful in vivo capabilities of bioluminescence with the subset-discriminating capabilities of fluorescence-activated cell sorting (88). Similarly, dual reporters that combine nuclear imaging techniques (e.g., PET) with fluorescence and/or bioluminescence for gene expression studies have also been developed (71, 73, 89, 90). Taking advantage of these new capabilities should allow pulmonary scientists a "view" on lung biology that could not be imaged—or even imagined—a few years ago.
The authors thank colleagues in the Washington University Molecular Imaging Center and the microPET Facility for their help and insights in completing work described herein. The authors also thank Dr. Mark Conradi for many helpful discussions related to MRI. This study was supported in part by NIH grants HL32815, P50 CA94056, NCI Small Animal Imaging Resource Program (SAIRP) grant CA83060. Received in original form June 8, 2003
This article has been cited by other articles:
|
||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||||
Top
Abstract
Introduction
scintigraphy and single photon emission computed tomography (SPECT), for example (
B activation and IkB kinase expression (
