/  Part III.7 – Cellular and Molecular Imaging in Translational Cardiovascular Research



Cellular and Molecular Imaging in Translational Cardiovascular Research

Johannes Riegler PhD and Joseph C Wu MD PhD

A. Introduction

Over the last decades, translational cardiovascular research has focused on elucidating molecular mechanisms underlying pathological processes and their alleviation via targeted intervention. These interventions cover a wide range in scale and type from small molecules to cells and mechanical devices. While cell culture models are instrumental for the study of molecular pathways, they cannot replicate the complexity of the disease on a tissue or organ level. In order to address this limitation, animal models have been developed with the aim of replicating human diseases. Information on the behavior of cells and molecular processes has traditionally been derived from histological examinations. However, histological examinations are time consuming, prone to sampling bias, and limited to examinations requiring tissue sampling ex vivo and in many cases sacrificing the animal. Cellular and molecular imaging methods have been developed to allow non-invasive assessment of cellular and molecular processes over time in animals and humans.

Cellular imaging in this context aims to visualize the distribution, movement, and behavior of cells in a living organism. Molecular imaging, on the other hand, visualizes processes that take place on a subcellular level such as the activation of a particular molecular pathway, the expression of a specific gene or the presence of certain proteins on the cell surface. Molecular imaging may overlap with cellular imaging, particularly if processes are imaged that lead to behavioral changes of cells. Although the ability to resolve single cells would be desirable, current technology limits the resolution to a few millimeters or tens of micrometers depending on the imaging modality. A variety of imaging modalities ranging from optical to radioactive and non-ionizing methods are available for molecular and cellular imaging. Major advantages and disadvantages of individual techniques are summarized in Table 1, and their basic principles as well as example applications are outlined below and illustrated in Fig. 1 and Fig. 2.

Table 1: Imaging modalities

table 1

Fields in blue indicate parameters specific to clinical systems if they are different from preclinical systems. Cost categories: low <50,000, medium <200,000, and high >500,000 USD.

figure 1Figure 1. Schematic overview of cellular and molecular imaging methods. Direct labeling of cells with fluorescent dyes, iron oxide nanoparticle, or radioisotopes is an easy option to label cells for cell tracking with fluorescence tomography (FMT), magnetic resonance imaging (MRI), or single photon computes tomography (SPECT) / positron emission tomography (PET), respectively. Direct labeling cannot assess cell viability or differentiation, which can be assessed with reporter gene imaging. For reporter gene imaging, a gene encoding a specific transporter, fluorescent protein, or enzyme is expressed. The sodium / iodine symporter (NIS) leads to active uptake of pertechnetate, which can be imaged with SPECT. In a similar fashion, the herpes simplex virus 1-thymidine kinase (HSV1-tk) phosphorylates a PET substrate, leading to its accumulation in cells. Luciferase oxidizes Luciferin and produces light during the process, which is captured with a bioluminescence imaging (BLI) system. Molecular imaging uses imaging tracers attached to small molecules, peptides, proteins, or antibodies that bind to specific targets on a cell.

figure 2Figure 2. Example images for cell tracking, reporter gene imaging, and molecular imaging with major imaging modalities. The first column shows bioluminescence imaging (BLI) of: (a) proliferation and tumor formation after embryonic stem cell injection into a mouse heart. Reproduced from (17) with permission from Wolters Kluwer Health; (b) survival of cardiac stem cells following myocardial injection. Reproduced from (18) with permission from Elsevier; and (c) in vivo expression of luciferase following minicircle delivery to rat hearts. Reproduced from (25) with permission from Wolters Kluwer Health. The second column shows example images from positron emission tomography (PET) of: (d) canine induced pluripotent stem cells injected into the myocardium of dogs. Reproduced from (71) with permission from The Journal of Biological Chemistry; (e) proliferation of simplex virus 1-thymidine kinase transfected embryonic stem cells injected into rat hearts (67); and (f) binding of radiolabeled vascular endothelial growth factor (VEGF) in an infarcted rat heart. Reproduced from (65) with permission from Wolters Kluwer Health. The third column shows magnetic resonance imaging (MRI) of: (g) iron oxide nanoparticle labeled cells in the myocardium of a dog after myocardial injection. Reproduced from (71) with permission from The Journal of Biological Chemistry; (h) MRI reporter gene imaging of transfected cells in a rat brain. Reproduced from (103) with permission from John Wiley and Sons; and (i) T2 map of annexin conjugated iron oxide nanoparticles targeting apoptotic cells in the infarcted myocardium of a mouse. Reproduced from (118) with permission from John Wiley and Sons.

B. Optical Imaging Methods

Optical imaging methods can be divided into two major groups. The first group uses the interaction of light from an external light source with molecules in the animal to generate an image. Frequently used techniques are fluorescence tomography (FMT) and optical coherence tomography (OCT). The second group records the light produced as a result of biochemical processes catalyzed by enzymes such as luciferase, which has been used for bioluminescence imaging (BLI). All optical imaging methods suffer from rapid signal loss with increasing tissue depth due to scattering and absorption, which limits their applicability to small animals. However, they offer high sensitivity and specificity, in particular BLI.

Fluorescence Tomography (FMT)

FMT is a diffuse optical tomography imaging technique that uses multiple light sources or a laser scanning over the object to excite fluorophores and a sensitive charge coupled device camera (CCD) to record the emitted fluorescence. Modern FMT systems correct for differences in fluorescent light propagation of tissues based on x-ray attenuation or normalization to transmitted excitation light, thereby improving image resolution and signal quantification from internal organs (1). Fluorophores with emission maxima in the near infrared wavelength band are typically used because living tissue absorbs less light in the near infrared range1. Image contrast is generated via direct labeling of cells or small molecules with organic fluorophores or fluorescent nanoparticles such as quantum dots. Alternatively, a reporter gene can be expressed that encodes a fluorescent protein or an enzyme that activates an administered substrate (2). Direct labeling and reporter gene imaging can be used for cellular and molecular imaging. Advantages of FMT are its good sensitivity and specificity, quantification capability, and low price. Low resolution, long imaging time, and limited clinical translatability are its major disadvantages.

Sosvonik et al illustrated FMT based tracking of macrophages labeled with magneto-fluorescent nanoparticles migrating to the infarcted mouse heart (3). The potential to image molecular processes such as the expression of tumor associated proteases, the activation of matrix metalloproteinases, or the catalytic activity of cathepsin-B in atherosclerotic plaques via FMT has been demonstrated in animal models (4-6). A possible application of FMT for the detection of breast cancer in humans after intravenous injection of a fluorescent contrast agent has been reported (7).

Optical Coherence Tomography (OCT)

OCT is a high resolution cross-sectional imaging technique mainly used for intravascular procedures similar to intravascular ultrasound. A light beam is split and one half is directed at the tissue (arterial wall). The reflected light is combined with the second half of the light beam (reference beam). Changes in the travel time of the light reflected by the tissue lead to interference, which is translated into image contrast (8). Near infrared light is used to increase tissue penetration, allowing an imaging depth of 2-3 mm and a spatial resolution of 10-15 µm. OCT systems are used for the assessment of vulnerable plaques and for monitoring the arterial wall after intravascular procedures such as stent placement (9-11). Cellular imaging has not been performed with OCT, but new developments – particularly the combination of OCT with near infrared fluorescence imaging, might allow cell imaging. Another modification of OCT measures the birefringence, the change in polarization of light due to the presence of aligned macromolecules such as collagen or actin, allowing direct molecular imaging. This technique has been used to characterize plaque collagen as well as normal, fibrous, calcified, and lipid rich plaques (12,13). Micron sized iron oxide nanoparticles bound to antibodies specific for endothelial markers VECAM-1 and PECAM-1 have demonstrated the molecular imaging potential of OCT (14).

Bioluminescence Imaging (BLI)

The ability of certain organisms to generate light from biochemical reactions has initiated research for its application as an imaging tool. Luciferase enzymes, which control such reactions, have been cloned from different organisms for use as genetic reporters (15). Luciferases emitting light at longer wavelengths (> 600 nm) are commonly used to improve tissue penetration. These are typically derived from North American firefly (Photinus pyralis), sea pansy (Renilla reniformis), or click beetle (Pyrophorus plagiophalamus). Expression of reporter genes can be controlled via constitutive promoters such as ubiquitin or tissue-specific promoters, or through recombination or administration of a chemical compound. For maximum sensitivity, strong viral promoters can be used to express high levels of luciferase in the cytoplasm16. Uptake of the substrate (D-Luciferin for firefly luciferase) into cells leads to its enzymatic conversion and light emission during the reaction. Animals are placed into a dark chamber where emitted light from different body parts is captured with a sensitive CCD camera (16). Since normal tissues do not express luciferases or other enzymes that convert D-Luciferin, there is no inherent background signal, leading to the high sensitivity of BLI. The combination of high sensitivity, low costs, flexibility of reporter genes to assess different cellular and molecular processes, and short imaging times have led to the widespread use of this technique. But BLI also has some drawbacks, such as low resolution, limited volumetric imaging capabilities, and its semi-quantitative nature. Tissue scattering and absorption limit the resolution and the ability for absolute quantification, as cells deeper in the body expressing the same amount of luciferase generate a weaker bioluminescence signal compared to cells close to the skin. Nevertheless, a suitable experimental design and rigorous validation can allow the quantification of signals from deep tissues. Depending on the experimental settings as few as 1000 cells can be detected if injected subcutaneously (17). Clinical translatability of BLI is limited, because the imaging depth is limited to a few centimeters and the target cells would need to be transfected with a reporter gene.

Although clinical translation of BLI is currently limited, cardiovascular research has significantly benefited from BLI. In particular the ability to assess cellular and molecular processes in living animals over time has improved our understanding of cell migration, immune rejection, and gene therapy. For instance, BLI has been used to demonstrate that, regardless of the cell type transplanted, less than 10% of the initially engrafted cells survive after the first three weeks (18-20) (see Fig. 2). Poor homing and accumulation in organs such as the lung, liver, and spleen has been observed with BLI after intravenous cell infusion (21,22). The ability to monitor cell survival over time with BLI was used to follow cell death after myocardial cell delivery. High levels of cell death around day 4 post delivery were correlated with T cell infiltration, indicating acute rejection (23). Long-term cell survival is further limited by the adaptive immune system responding to the foreign antigens. Pearl et al used BLI to demonstrate that a mixture of costimulatory blockers (anti-CD40 ligand, CTLA4 antibodies, and anti-lymphocyte function associated antigen) could be used to prevent the adaptive immune rejection of human embryonic stem cells and pluripotent stem cells injected in mice (24). Although most of the aforementioned examples can be classified as cellular imaging, BLI is also well suited for molecular imaging, as has been demonstrated for the assessment of gene therapies (25-27) (see Fig. 2) or cell differentiation (28,29).

C. Nuclear Imaging Methods

Nuclear imaging methods use the radiation released from unstable elements to generate an image. Depending on the type of radiation used, two techniques are the most common. Single photon emission computed tomography (SPECT) detects γ-radiation directly, while positron emission tomography (PET) detects pairs of gamma rays that are produced through the annihilation of a positron. Both imaging methods allow tomographic imaging without depth limitations and offer high sensitivity. This has led to their widespread use in clinical diagnosis. Nuclear imaging methods are particular well suited for whole body screening and bio-distribution studies due to the fast acquisition of tomographic images. Computed tomography (CT) is another imaging modality based on the use of ionizing radiation and will hence be discussed together with PET and SPECT.

Single Photon Emission Computed Tomography (SPECT)

SPECT imaging systems consist of a set of gamma cameras that are arranged around the subject. These cameras are equipped with collimators (metal plates), which limit the sensitivity of the camera to γ-rays from a small region of the subject. Cameras are rotated slowly to receive γ-rays from the whole imaging volume. This process generates a set of 2D images (projections), which can be combined to make a 3D imaging volume using a mathematical process (e.g., back projection) (30). Radioisotopes frequently used for SPECT imaging include Technetium-99, Indium-111, and Iodine-123 with respective half live times of 6, 67, and 13 hours, limiting the maximum tracking time for a single administration to a few days. Clinical SPECT systems achieve a resolution of 8-12 mm, whereas dedicated preclinical systems can reach less than 1 mm. In order to improve the spatial localization of the SPECT signal, imaging systems are typically combined with computed tomography (CT) or magnetic resonance imaging (MRI) to provide high resolution reference images of the tissue (30). These combined systems offer improved quantification capabilities, because γ-ray attenuation due to tissue density differences can be accounted for. SPECT imaging offers high sensitivity, good image resolution, volumetric imaging, and good clinical translatability at a medium price. Exposure to ionizing radiation and the short half-life of radioisotopes are its major disadvantages.

Direct labeling of cells with lipophilic tracers such as Indium-111-oxine or technetium-99-hexamethylprophylene-amine-oxine (HMPAO) has been used frequently to image cell distribution kinetics in humans and animals (31-35). Kraitchman et al demonstrated that In111 labeled mesenchymal stem cells (MSCs) initially accumulate in the lung following intravenous infusion in dogs (35). This is followed by redistribution of MSCs within 24-48 hours, leading to increased accumulation in the liver, kidney, and infarcted areas of the heart (see Fig. 2). However, maximum accumulation of MSCs in the heart was only 4.6% of the administered dose. In agreement with this result, Hou et al showed that one hour after myocardial injection, coronary artery infusion or retrograde venous delivery only 11.3, 2.6, and 3.2% of delivered cells respectively, remained in the heart (33).

SPECT has also been used for molecular and reporter gene imaging. A wide range of small molecules (36-38), antibodies (39-42), and proteins (43) labeled with radioisotopes have been used to assess the expression of cell surface receptors and intracellular proteins. Hofstra et al demonstrated the ability to image apoptosis following myocardial infarction in humans with radiolabelled annexin (43). SPECT can also be used to assess infarct size (42) and tissue perfusion (44). Differences in the energy signature of radioisotopes can be used to perform multiplexed imaging such as the assessment of myocardial perfusion and distribution of labeled cells in one imaging experiment. Beeres et al illustrated that bone marrow cell delivery led to a reduction in the number of segments with stress inducible ischemia in humans (45). SPECT can be used for reporter gene imaging to assess cell viability and gene expression over time. The sodium / iodide symporter (NIS) is frequently used as a reporter gene. NIS is normally expressed in the thyroid, salivary glands, stomach, and choroids plexus. Cells transfected with NIS import iodide or pertechnetate (Technetium-99m) from body fluids, causing local accumulation and signal enhancement after the blood pool has been cleared of the tracer. Reporter gene imaging has been used frequently to track cells after myocardial delivery. Templin et al imaged viability and localization of NIS-transfected human induced pluripotent stem cells (hIPSCs) in pig hearts up to 15 weeks post cell delivery via SPECT (46). Reporter gene imaging was also used to monitor gene expression up to 12 days in rat hearts after adenoviral gene delivery (47) (see Fig. 2).

Positron Emission Tomography (PET)

PET imaging uses positron-emitting radionuclides to generate image contrast. Emitted positrons travel a short distance until they interact with an electron, leading to the annihilation of both positron and electron as well as the release of two γ-rays at nearly 180 degrees from each other. These γ-rays are detected by a circular detector array placed around the subject. Simultaneous detection of both rays is used to discard background scattering and reduce noise48. Detected events can be arranged into projections and reconstructed into a 3D volume similar to SPECT or CT. Frequently used radioisotopes are Flourine-18 (18F), Carbon-11 (11C) and Copper-64 (64Cu) with respective half-lives of 110 minutes, 20 minutes, and 12.7 hours (48). Their short half-lives limit the distance radioisotopes can be transported and require access to a cyclotron for their production. This increases the cost of PET imaging and has limited its clinical use primarily to 18F, which can be transported more easily. 18F and 11C are integrated into small organic molecules such as the glucose analog fluorodeoxyglucose (18F-FDG), carbon-11-choline, or carbon-11-acetate before intravenous administration (49,50). Depending on their structure, radiotracers are cleared from the blood and accumulate in target tissues. For example, 18F-FDG is taken up by metabolically active cells, thus highlighting tissues with high metabolic activity such as the brain (51) or tumors (52,53). Similar to SPECT, PET is frequently combined with CT or MRI to improve the spatial localization of the radioactive signal. PET imaging has a slightly higher sensitivity compared to SPECT, and achieves a similar resolution of less than 10 mm on clinical and 1-2 mm on preclinical systems (54). Other advantages include the lack of depth limitation, volumetric imaging, easy quantification, clinical translatability, and the possibility to measure metabolic rates with 11C compounds. Major disadvantages of PET imaging are mutation risks due to ionizing radiation and high cost (55).

Cellular imaging via direct labeling with 18F-FDG has been used frequently to assess cell distribution and migration in animals and humans (56-59). Hofmann et al demonstrated that 75 minutes after intracoronary infusion of labeled cells in humans, less than 2.6 % of the cells were found in the myocardium (57). They also found that CD34-enriched cell populations might lead to higher engraftment and that administered cells generally engrafted in the border zone of the infarct. 18F-FDG can also be used to assess the viability of the host myocardium (60). However, to combine myocardial viability assessment and PET cell tracking, radiotracers with sufficient half-live difference are necessary since multiplexing is not possible with PET. Several cardiovascular molecular imaging applications have been reported, including reporter gene imaging. Nahrndorf et al used 18F labeled peptides specific for VCAM-1 to image atherosclerotic plaques in apolipoprotein deficient mice61. The same group has demonstrated the ability to track macrophages labeled with iron oxide nanoparticles conjugated with 64Cu chelate in the same mouse model (62). Other applications have also been reported, such as angiogenesis assessment over time using radioactively labeled peptides binding to αvβ3-integrin (63,64) and imaging of vascular endothelial growth factor receptor expression in the infarcted rat heart (65) (see Fig. 2).

Previously mentioned advantages of reporter gene imaging, such as the ability to track living cells over long periods of time, to differentiate between living and dead cells, and to monitor gene expression, are equally true for PET reporter gene imaging. Common reporter genes include the herpes simplex virus 1-thymidine kinase (HSV 1-tk) and its mutated form HSV1-sr39tk. HSV 1-tk phosphorylates its substrate 9-4-18F-fluoro-3-hydroxymethylbutyl-guanine (18F-FHBG) and similar agents, leading to their accumulation in transfected cells. By contrast, the sodium / iodide symporter (NIS) can also be used with 124I for PET. The ability of these reporter genes to track cell survival and growth has been compared with BLI in rats (66) and has been used for cardiac cell delivery in animal models (67-71). Cao et al showed that a combination of 18F-FDG to visualize viable myocardium und 18F-FHBG for reporter gene imaging could be used to demonstrate the location of cells in the myocardial wall and potential tumors from injected embryonic stem cells (67) (see Fig. 2). An additional advantage of HSV 1-tk is its potential use as a suicide gene, because it can convert ganciclovir to its monophosphate form, which blocks DNA synthesis after conversion to a triphosphate (72).

Computed Tomography (CT)

CT uses a focused X-ray beam directed at the subject with a detector on the opposite side to generate images based on beam attenuation. X-ray beam and detector are rotated around the subject, and the subject is moved slowly along the anterior-posterior axis to generate a set of 2D projections that are reconstructed to a 3D volume (via back projection) (73). Image contrast is generated by x-ray attenuation differences caused by tissue density differences. CT scanners achieve resolutions of less than 1 mm on clinical systems and less than 0.05 mm on preclinical systems. Additional advantages of CT imaging are its relatively low cost, fast imaging time, and ease of use. Its major drawbacks are the poor soft tissue contrast and the mutation risk due to the use of ionizing radiation (74). Potential applications of CT for cellular and molecular imaging have been reported recently (75). Jensen et al showed that the small angle X-ray scattering CT can be used to map the myelin distribution in a rat brain (75). Other cellular and molecular imaging applications are based primarily on gold nanoparticles used for cell tracking (76,77) and enhanced cancer detection (78). Nevertheless, the sensitivity of this technique is lower compared to SPECT, PET and MRI.

D. Non-ionizing Imaging Methods

The two most important modalities in this group besides optical methods are ultrasound (US) imaging and magnetic resonance imaging (MRI) / magnetic resonance spectroscopy (MRS). Ultrasound is a fast and cheap technique that is primarily used for the assessment of cardiac and vascular functions, but new contrast agents may enable US imaging of cells and molecular targets. MRI offers unique soft tissue contrast and high spatial resolution, which has led to its widespread use in clinical diagnostics. Cellular and molecular MRI has been demonstrated with suitable contrast agents and might soon be used in clinical diagnostics. MRS on the other hand, is an inherently molecular technique common in bimolecular research and might find more clinical applications with technical improvements.

Ultrasound (US)

US imaging devices use piezoelectric transducers or transducer arrays to generate ultrasound waves (2-18 MHz). These sound waves travel through the body, where they are partially reflected at tissue interfaces with different densities, generating an echo. Echoes are recorded with a receiver that determines the place where the echo originated by the time it took to receive it (79). Depending on the wavelength used, high resolution images (∼ 150 µm) with low penetration depth (1-2 cm) or low resolution images (1-3 mm) with higher penetration (15-20 cm) can be acquired with high or low wavelength, respectively (79). Different imaging modes on US systems allow the generation of 2D images and, on modern systems, 3D images of the anatomy as well as flow and motion measurements in the Doppler mode (80). Major advantages of US imaging are fast real-time imaging, good contrast between fluid and soft tissues, easy blood flow measurements, and cheap compact imaging systems. Major drawbacks of US imaging are limited depth penetration, limited tomographic capability, low soft tissue contrast, and variable image quality depending on operator ability. Dedicated high frequency transducers for small animal imaging and intravascular ultrasound are available that can achieve a resolution of less than 50 µm.

Molecular and cellular US imaging uses targeted contrast agents such as microbubbles (81), liposomes (82), or solid nanoparticles (83). Microbubbles consist of a gaseous core with an organic shell preventing aggregation and are typically 1-10 µm in diameter. Due to their size, only targets present on the vascular wall can be detected. Nanobubbles, liposomes, and nanoparticles allow diffusion or extravasation into tissue. Since they are difficult to detect, most of the research has focused on microbubbles. Lindner et al demonstrated the ability to monitor inflammation with microbubbles attached to antibodies specific for P-selectin in a mouse model of kidney injury (84). A similar approach with an analogues contrast agent was used to image increased P-selectin presence in rat hearts following ischemia / reperfusion injury (85). Kuliszewski et al indicated that endothelial progenitor cells, which had been transfected with a reporter gene (H-2Kk) and subcutaneously injected in a rat (106 cells + Matrigel), could be visualized 7 days after implantation with microbubbles specific to H-2Kk (86).

Magnetic Resonance Imaging (MRI) / Magnetic Resonance Spectroscopy (MRS)

For MRI, the subject of interest is placed in a strong magnetic field that causes the alignment of proton magnetic moments (spins) parallel or antiparallel to the external magnetic field following a Boltzmann distribution. Spins can be rotated away from this equilibrium through the application of an additional magnetic field generated by a radio frequency (RF) pulse. Following this distortion, the spins return to their equilibrium state. A radio frequency coil, in the simplest case as a wire loop, placed next to the subject can pick up a decaying oscillating current induced by the rotating spins (87). Spins can return to their equilibrium state via two mechanisms, spin-lattice and spin-spin relaxation, which take place simultaneously and are defined by their characteristic time constants. The spin-lattice relaxation constant or T1 relaxation time is influenced by the mobility and interaction of protons. Spin-spin relaxation times or T2 relaxation times are dependent on the magnetic interaction between protons. In MRI, a series of exactly timed magnetic field gradients and RF pulses (MR imaging sequences) are used to record the signal from decaying spins of each volumetric element (voxel) making up the image volume (87). A mathematical operation (Fourier transformation) can be used to convert this data into images. Depending on the parameters used for the imaging sequence, images with different weighting towards T1, T2, or proton density can be generated (87). Inherent differences for these parameters between different tissues are responsible for the excellent soft tissue contrast of MRI. MRI offers high-resolution tomographic images (<1 mm on clinical systems and <100 µm on preclinical systems) without depth limitation or the use of ionizing radiation. Image contrast can be quantitative and reflects the molecular properties of the tissue. These advantages have led to the widespread use of MRI in clinical practice and preclinical research facilitating translational research. Major disadvantages are the high cost and complexity of MRI, lower temporal resolution compared to other modalities (depending on imaging sequence), lower sensitivity relative nuclear methods, and incompatibility with ferromagnetic implants.

Image contrast and detection sensitivities for certain pathologies can be further improved through the use of contrast agents. These are classified into two major groups: T1 and T2 contrast agents, depending on their mode of action. T1 contrast agents such as Gadolinium or Manganese reduce the T1 relaxation time and appear bright on T1 weighted images, providing “positive contrast”. Conversely, T2 contrast agents such as iron oxide nanoparticles appear dark on T2 weighted images, generating hypointensities or “negative contrast”. Gradient echo and related imaging sequences are particularly sensitive to T2 contrast agents, and allow the detection of single cells in vivo if cells are labeled with a sufficient amount of iron oxide typically via the use of micron sized nanoparticles (88). Micron sized particles (0.5-5 µm) provide maximum sensitivity, but many of these particles are not biodegradable. However, biodegradable iron oxide nanoparticles (20-200 nm) are available, including some FDA-approved MRI contrast agents (Feridex, Combidex and Sinerem). Different protocols are available for labeling cells with these particles, ranging from simple incubation to the use of transfection agents (89,90). Several animal studies have demonstrated the potential of MRI based cell tracking (71,91-95) (see Fig. 2). Hill et al showed that as few as 105 MSCs labeled with iron oxide nanoparticles injected into an infarcted pig heart could be imaged for up to 3 weeks with a clinical MRI scanner (93). A similar study by Carr et al indicated that a range of cell types injected into infarcted rat hearts could be tracked over time, but did not lead to improved cardiac function (95). However, one of the problems of MRI-based cell tracking is the inability to differentiate dead from living cells. While hypo-intensities might persist for weeks, cells can die and macrophages tend to take up released nanoparticles, leading to false positive detection of cells (96). Clinical cell tracking has been demonstrated in several pilot studies with the off label use of MRI contrast agents (97-99). The high resolution of MRI enabled de Vries et al to identify cell injections that did not reach the target lymph node, which was not possible with nuclear imaging (97). Furthermore, cell migration to neighboring lymph nodes could be visualized and a detection limit of 15,000 cells was found for their clinical setup. Cell labeling and tracking with T1 contrast agents is also feasible but generally less sensitive, and the toxicity of these agents remains a concern (100). To address the problem of distinguishing between live and dead cells, reporter gene approaches for MRI have been explored (101-103) (see Fig. 2). Unfortunately these techniques are unreliable and have not been demonstrated for cardiac applications.

MRI is almost exclusively used to image hydrogen nuclei. However, other nuclei with unpaired protons and/or neutrons such as Fluorine-19, Phosphors-31, Carbon-13, or Helium-3 can be image as well. Low molar concentrations of these nuclei lead to weaker signals. However, with elements such as Fluorine, there is essentially no background signal from the body enabling high specificity. The limitation is the lower sensitivity for the detection of Fluorine-labeled entities (104,105). This specificity was explored by Flögel et al to track macrophages labeled with perfluorocarbons infiltrating the border region of the infarcted myocardium in mice (105). As few as 6000 cells labeled with perfluorocarbons could be detected with a preclinical MRI system (104).

Molecular imaging with MRI has been performed with both T1 and T2 contrast agents. Preferential distribution (106,107), specific binding (108-110), and enzyme activated (111) T1 contrast agents have been used for that. Amirbekian et al used macrophage scavenger receptor targeted immunomicelles containing Gadolinium to visualize macrophages in atherosclerotic plaques in ApoE-/- mice (110). A twofold enhancement in macrophage-rich plaques could be detected one hour after intravenous contrast administration. Enzyme activated contrast agents can offer increased sensitivity via contrast agent accumulation. This approach was explored by Nahrendorf et al with a Gadolinium-based contrast agent, which, after activation through myeloperoxidase, polymerizes and accumulates in areas of inflammation111. Accordingly, T1 weighted MRI could be used to assess the inflammatory response following ischemia reperfusion injury. As for cell labeling, iron oxide nanoparticles are most frequently used for T2 weighted molecular imaging. This is typically done with nanoparticles bound to antibodies (112-115) or small peptides (116-118) binding to the target of interest. Depending on the size of the particles used, primary intravascular or some extravascular targets (using small particles) can be detected. Small iron oxide nanoparticles (9 nm), which can reach extravascular targets, attached to Herceptin antibodies were used by Huh et al to detect HER2 expressing cancer cells in a mouse model (112). Slightly bigger nanoparticles (50 nm) attached to Annexin were used by Sosnovik et al to visualize areas of apoptotic cell death following ischemia reperfusion injury in mice (118) (see Fig. 2).

MRI uses the intensity of the characteristic resonance frequency of protons (or other nuclei) to generate an image. However, protons in a different chemical environment have a resonance frequency or several frequencies that are offset from the characteristic frequency. MRS acquires a spectrum that displays peaks for different resonance frequencies, which are directly correlated to the concentration of these substances (87). Accordingly, MRS can be used to measure the concentration of chemical substances in vivo. For example, proton spectroscopy can detect the triglycerides deposited in obese rat hearts and humans (119). Phosphorous spectra can be acquired from the heart to measure the concentrations of ATP, phosphocreatine, and inorganic phosphate as indicators of the metabolic state of the tissue (120,121). The low signal-to-noise ratio of MRS and consequently slow acquisition is the main disadvantage of MRS. However, new developments, in particular the use of hyperpolarization (122,123), parallel imaging (124), and sparse MRI (125), have led to tremendous signal-to- noise ratio improvements, which my translate to increased clinical use.

E. Conclusions

A wide range of methods for preclinical cellular and molecular imaging is available. Unfortunately none of them is ideal and some of them (BLI, OCT, FMT) are not yet clinically translatable. Thus, a combination of different modalities with complementary qualities may be necessary to answer specific questions. Careful considerations are necessary to define the requirements for a study and match them with the available technology.


This work was supported by the Austrian Science Fund (FWF, Erwin Schrödinger Fellowship) (JR) and NIH R01 EB009689, R01 HL093172, and R01 HL095571 (JCW).



  1. Frangioni, J.V. In vivo near-infrared fluorescence imaging. Curr Opin Chem Biol 7, 626-634 (2003).
  2. Ntziachristos, V., Bremer, C. & Weissleder, R. Fluorescence imaging with near-infrared light: new technological advances that enable in vivo molecular imaging. Eur Radiol 13, 195-208 (2003).
  3. Sosnovik, D.E., et al. Fluorescence tomography and magnetic resonance imaging of myocardial macrophage infiltration in infarcted myocardium in vivo. Circulation 115, 1384-1391 (2007).
  4. Choi, Y., Weissleder, R. & Tung, C.H. Selective antitumor effect of novel protease-mediated photodynamic agent. Cancer Res 66, 7225-7229 (2006).
  5. Chen, J., et al. Near-infrared fluorescent imaging of matrix metalloproteinase activity after myocardial infarction. Circulation 111, 1800-1805 (2005).
  6. Chen, J., et al. In vivo imaging of proteolytic activity in atherosclerosis. Circulation 105, 2766-2771 (2002).
  7. Corlu, A., et al. Three-dimensional in vivo fluorescence diffuse optical tomography of breast cancer in humans. Opt Express 15, 6696-6716 (2007).
  8. Fujimoto, J.G., et al. Optical biopsy and imaging using optical coherence tomography. Nat Med 1, 970-972 (1995).
  9. Takano, M., Jang, I.K. & Mizuno, K. Neointimal proliferation around malapposed struts of a sirolimus-eluting stent: optical coherence tomography findings. Eur Heart J 27, 1763 (2006).
  10. Tearney, G.J., Jang, I.K. & Bouma, B.E. Optical coherence tomography for imaging the vulnerable plaque. J Biomed Opt 11, 021002 (2006).
  11. Toutouzas, K., Vaina, S., Riga, M.I. & Stefanadis, C. Evaluation of dissection after coronary stent implantation by intravascular optical coherence tomography. Clin Cardiol 32, E47-48 (2009).
  12. Giattina, S.D., et al. Assessment of coronary plaque collagen with polarization sensitive optical coherence tomography (PS-OCT). Int J Cardiol 107, 400-409 (2006).
  13. Kuo, W.C., et al. Polarization-sensitive optical coherence tomography for imaging human atherosclerosis. Appl Opt 46, 2520-2527 (2007).
  14. Jefferson, A., et al. Molecular imaging with optical coherence tomography using ligand-conjugated microparticles that detect activated endothelial cells: rational design through target quantification. Atherosclerosis 219, 579-587 (2011).
  15. Hastings, J.W. Chemistries and colors of bioluminescent reactions: a review. Gene 173, 5-11 (1996).
  16. Contag, C.H., et al. Visualizing gene expression in living mammals using a bioluminescent reporter. Photochem Photobiol 66, 523-531 (1997).
  17. Cao, F., et al. Spatial and temporal kinetics of teratoma formation from murine embryonic stem cell transplantation. Stem Cells Dev 16, 883-891 (2007). Image reprinted with permission from Mary Ann Liebert, Inc.
  18. Li, Z., et al. Imaging survival and function of transplanted cardiac resident stem cells. J Am Coll Cardiol 53, 1229-1240 (2009). Image reprinted with permission from Elsevier.
  19. Li, Z., et al. Differentiation, survival, and function of embryonic stem cell derived endothelial cells for ischemic heart disease. Circulation 116, I46-54 (2007).
  20. van der Bogt, K.E., et al. Comparison of different adult stem cell types for treatment of myocardial ischemia. Circulation 118, S121-129 (2008).
  21. Schrepfer, S., et al. Stem cell transplantation: the lung barrier. Transplant Proc 39, 573-576 (2007).
  22. Sheikh, A.Y., et al. Molecular imaging of bone marrow mononuclear cell homing and engraftment in ischemic myocardium. Stem Cells 25, 2677-2684 (2007).
  23. Tanaka, M., et al. In vivo visualization of cardiac allograft rejection and trafficking passenger leukocytes using bioluminescence imaging. Circulation 112, I105-110 (2005).
  24. Pearl, J.I., et al. Short-term immunosuppression promotes engraftment of embryonic and induced pluripotent stem cells. Cell Stem Cell 8, 309-317 (2011). Image reprinted with permission from Elsevier.
  25. Huang, M., et al. Novel minicircle vector for gene therapy in murine myocardial infarction. Circulation 120, S230-237 (2009).
  26. Rehemtulla, A., et al. Molecular imaging of gene expression and efficacy following adenoviral-mediated brain tumor gene therapy. Mol Imaging 1, 43-55 (2002).
  27. Wu, J.C., Sundaresan, G., Iyer, M. & Gambhir, S.S. Noninvasive optical imaging of firefly luciferase reporter gene expression in skeletal muscles of living mice. Mol Ther 4, 297-306 (2001).
  28. Iris, B., et al. Molecular imaging of the skeleton: quantitative real-time bioluminescence monitoring gene expression in bone repair and development. J Bone Miner Res 18, 570-578 (2003).
  29. Xie, X., et al. Imaging of STAT3 signaling pathway during mouse embryonic stem cell differentiation. Stem Cells Dev 18, 205-214 (2009).
  30. Gullberg, G.T., Reutter, B.W., Sitek, A., Maltz, J.S. & Budinger, T.F. Dynamic single photon emission computed tomography–basic principles and cardiac applications. Phys Med Biol 55, R111-191 (2010).
  31. Aicher, A., et al. Assessment of the tissue distribution of transplanted human endothelial progenitor cells by radioactive labeling. Circulation 107, 2134-2139 (2003).
  32. Barbosa da Fonseca, L.M., et al. Biodistribution of bone marrow mononuclear cells in chronic chagasic cardiomyopathy after intracoronary injection. Int J Cardiol 149, 310-314 (2011).
  33. Hou, D., et al. Radiolabeled cell distribution after intramyocardial, intracoronary, and interstitial retrograde coronary venous delivery: implications for current clinical trials. Circulation 112, I150-156 (2005).
  34. Kollaros, N., et al. Bone marrow stem cell adherence into old anterior myocardial infarction: a scintigraphic study using Tl-201 and Tc-99m-HMPAO. Ann Nucl Med 26, 228-233 (2012).
  35. Kraitchman, D.L., et al. Dynamic imaging of allogeneic mesenchymal stem cells trafficking to myocardial infarction. Circulation 112, 1451-1461 (2005).
  36. Meegalla, S.K., et al. Synthesis and characterization of technetium-99m-labeled tropanes as dopamine transporter-imaging agents. J Med Chem 40, 9-17 (1997).
  37. Wagner, S., et al. Molecular imaging of matrix metalloproteinases in vivo using small molecule inhibitors for SPECT and PET. Curr Med Chem 13, 2819-2838 (2006).
  38. Su, H., et al. Noninvasive targeted imaging of matrix metalloproteinase activation in a murine model of postinfarction remodeling. Circulation 112, 3157-3167 (2005).
  39. Backer, M.V., et al. Molecular imaging of VEGF receptors in angiogenic vasculature with single-chain VEGF-based probes. Nat Med 13, 504-509 (2007).
  40. Divgi, C.R., et al. Phase I and imaging trial of indium 111-labeled anti-epidermal growth factor receptor monoclonal antibody 225 in patients with squamous cell lung carcinoma. J Natl Cancer Inst 83, 97-104 (1991).
  41. Yasuda, T., et al. Indium 111-monoclonal antimyosin antibody imaging in the diagnosis of acute myocarditis. Circulation 76, 306-311 (1987).
  42. Gibbons, R.J., Miller, T.D. & Christian, T.F. Infarct size measured by single photon emission computed tomographic imaging with (99m)Tc-sestamibi: A measure of the efficacy of therapy in acute myocardial infarction. Circulation 101, 101-108 (2000).
  43. Hofstra, L., et al. Visualisation of cell death in vivo in patients with acute myocardial infarction. Lancet 356, 209-212 (2000).
  44. Germano, G., et al. Automatic quantification of ejection fraction from gated myocardial perfusion SPECT. J Nucl Med 36, 2138-2147 (1995).
  45. Beeres, S.L., et al. Sustained effect of autologous bone marrow mononuclear cell injection in patients with refractory angina pectoris and chronic myocardial ischemia: twelve-month follow-up results. Am Heart J 152, 684 e611-686 (2006).
  46. Templin, C., et al. Transplantation and tracking of human-induced pluripotent stem cells in a pig model of myocardial infarction: assessment of cell survival, engraftment, and distribution by hybrid single photon emission computed tomography/computed tomography of sodium iodide symporter transgene expression. Circulation 126, 430-439 (2012).
  47. Miyagawa, M., et al. Cardiac reporter gene imaging using the human sodium/iodide symporter gene. Cardiovasc Res 65, 195-202 (2005).
  48. Cherry, S.R. Fundamentals of positron emission tomography and applications in preclinical drug development. J Clin Pharmacol 41, 482-491 (2001).
  49. Hamacher, K., Coenen, H.H. & Stocklin, G. Efficient stereospecific synthesis of no-carrier-added 2-[18F]-fluoro-2-deoxy-D-glucose using aminopolyether supported nucleophilic substitution. J Nucl Med 27, 235-238 (1986).
  50. Hara, T., Kosaka, N. & Kishi, H. PET imaging of prostate cancer using carbon-11-choline. J Nucl Med 39, 990-995 (1998).
  51. Chugani, H.T., Phelps, M.E. & Mazziotta, J.C. Positron emission tomography study of human brain functional development. Ann Neurol 22, 487-497 (1987).
  52. Di Chiro, G., et al. Glucose utilization of cerebral gliomas measured by [18F] fluorodeoxyglucose and positron emission tomography. Neurology 32, 1323-1329 (1982).
  53. Pieterman, R.M., et al. Preoperative staging of non-small-cell lung cancer with positron-emission tomography. N Engl J Med 343, 254-261 (2000).
  54. Gambhir, S.S., et al. A tabulated summary of the FDG PET literature. J Nucl Med 42, 1S-93S (2001).
  55. Brix, G., et al. Radiation exposure of patients undergoing whole-body dual-modality 18F-FDG PET/CT examinations. J Nucl Med 46, 608-613 (2005).
  56. Dedobbeleer, C., et al. Myocardial homing and coronary endothelial function after autologous blood CD34+ progenitor cells intracoronary injection in the chronic phase of myocardial infarction. J Cardiovasc Pharmacol 53, 480-485 (2009).
  57. Hofmann, M., et al. Monitoring of bone marrow cell homing into the infarcted human myocardium. Circulation 111, 2198-2202 (2005).
  58. Kang, W.J., et al. Tissue distribution of 18F-FDG-labeled peripheral hematopoietic stem cells after intracoronary administration in patients with myocardial infarction. J Nucl Med 47, 1295-1301 (2006).
  59. Wang, J., et al. Human CD34+ cells in experimental myocardial infarction: long-term survival, sustained functional improvement, and mechanism of action. Circ Res 106, 1904-1911 (2010).
  60. Klein, C., et al. Assessment of myocardial viability with contrast-enhanced magnetic resonance imaging: comparison with positron emission tomography. Circulation 105, 162-167 (2002).
  61. Nahrendorf, M., et al. 18F-4V for PET-CT imaging of VCAM-1 expression in atherosclerosis. JACC Cardiovasc Imaging 2, 1213-1222 (2009).
  62. Nahrendorf, M., et al. Nanoparticle PET-CT imaging of macrophages in inflammatory atherosclerosis. Circulation 117, 379-387 (2008).
  63. Chen, X., et al. 18F-labeled RGD peptide: initial evaluation for imaging brain tumor angiogenesis. Nucl Med Biol 31, 179-189 (2004).
  64. Li, Z.B., et al. (64)Cu-labeled tetrameric and octameric RGD peptides for small-animal PET of tumor alpha(v)beta(3) integrin expression. J Nucl Med 48, 1162-1171 (2007).
  65. Rodriguez-Porcel, M., et al. Imaging of VEGF receptor in a rat myocardial infarction model using PET. J Nucl Med 49, 667-673 (2008). Image reprinted with permission from Wolters Kluwer Health and the Society of Nuclear Medicine, Inc.
  66. Wu, J.C., et al. Molecular imaging of cardiac cell transplantation in living animals using optical bioluminescence and positron emission tomography. Circulation 108, 1302-1305 (2003).
  67. Cao, F., et al. In vivo visualization of embryonic stem cell survival, proliferation, and migration after cardiac delivery. Circulation 113, 1005-1014 (2006). Image reprinted with permission from Wolters Kluwer Health.
  68. Qiao, H., et al. Death and proliferation time course of stem cells transplanted in the myocardium. Mol Imaging Biol 11, 408-414 (2009).
  69. Terrovitis, J., et al. Ectopic expression of the sodium-iodide symporter enables imaging of transplanted cardiac stem cells in vivo by single-photon emission computed tomography or positron emission tomography. J Am Coll Cardiol 52, 1652-1660 (2008).
  70. Gyongyosi, M., et al. Serial noninvasive in vivo positron emission tomographic tracking of percutaneously intramyocardially injected autologous porcine mesenchymal stem cells modified for transgene reporter gene expression. Circ Cardiovasc Imaging 1, 94-103 (2008).
  71. Lee, A.S., et al. Preclinical derivation and imaging of autologously transplanted canine induced pluripotent stem cells. J Biol Chem 286, 32697-32704 (2011). Images reprinted with permission from The American Society for Biochemistry and Molecular Biology.
  72. Cao, F., et al. Molecular imaging of embryonic stem cell misbehavior and suicide gene ablation. Cloning Stem Cells 9, 107-117 (2007).
  73. Goldman, L.W. Principles of CT and CT technology. J Nucl Med Technol 35, 115-128; quiz 129-130 (2007).
  74. Brenner, D.J. & Hall, E.J. Computed tomography–an increasing source of radiation exposure. N Engl J Med 357, 2277-2284 (2007).
  75. Jensen, T.H., et al. Molecular X-ray computed tomography of myelin in a rat brain. Neuroimage 57, 124-129 (2011).
  76. Menk, R.H., et al. Gold nanoparticle labeling of cells is a sensitive method to investigate cell distribution and migration in animal models of human disease. Nanomedicine 7, 647-654 (2011).
  77. Arifin, D.R., et al. Trimodal gadolinium-gold microcapsules containing pancreatic islet cells restore normoglycemia in diabetic mice and can be tracked by using US, CT, and positive-contrast MR imaging. Radiology 260, 790-798 (2011).
  78. Wang, H., et al. Computed tomography imaging of cancer cells using acetylated dendrimer-entrapped gold nanoparticles. Biomaterials 32, 2979-2988 (2011).
  79. Aldrich, J.E. Basic physics of ultrasound imaging. Crit Care Med 35, S131-137 (2007).
  80. Coatney, R.W. Ultrasound imaging: principles and applications in rodent research. ILAR J 42, 233-247 (2001).
  81. Klibanov, A.L. Preparation of targeted microbubbles: ultrasound contrast agents for molecular imaging. Med Biol Eng Comput 47, 875-882 (2009).
  82. Hamilton, A.J., et al. Intravascular ultrasound molecular imaging of atheroma components in vivo. J Am Coll Cardiol 43, 453-460 (2004).
  83. Nolte, I., et al. Iron particles enhance visualization of experimental gliomas with high-resolution sonography. AJNR Am J Neuroradiol 26, 1469-1474 (2005).
  84. Lindner, J.R., et al. Ultrasound assessment of inflammation and renal tissue injury with microbubbles targeted to P-selectin. Circulation 104, 2107-2112 (2001).
  85. Villanueva, F.S., et al. Myocardial ischemic memory imaging with molecular echocardiography. Circulation 115, 345-352 (2007).
  86. Kuliszewski, M.A., et al. Molecular imaging of endothelial progenitor cell engraftment using contrast-enhanced ultrasound and targeted microbubbles. Cardiovasc Res 83, 653-662 (2009).
  87. Haacke, E.M. Magnetic resonance imaging : physical principles and sequence design, (Wiley, New York, 1999).
  88. Shapiro, E.M., Sharer, K., Skrtic, S. & Koretsky, A.P. In vivo detection of single cells by MRI. Magn Reson Med 55, 242-249 (2006).
  89. Hinds, K.A., et al. Highly efficient endosomal labeling of progenitor and stem cells with large magnetic particles allows magnetic resonance imaging of single cells. Blood 102, 867-872 (2003).
  90. Arbab, A.S., et al. Efficient magnetic cell labeling with protamine sulfate complexed to ferumoxides for cellular MRI. Blood 104, 1217-1223 (2004).
  91. Kraitchman, D.L., et al. In vivo magnetic resonance imaging of mesenchymal stem cells in myocardial infarction. Circulation 107, 2290-2293 (2003).
  92. Bulte, J.W., et al. Magnetodendrimers allow endosomal magnetic labeling and in vivo tracking of stem cells. Nat Biotechnol 19, 1141-1147 (2001).
  93. Hill, J.M., et al. Serial cardiac magnetic resonance imaging of injected mesenchymal stem cells. Circulation 108, 1009-1014 (2003).
  94. Himes, N., et al. In vivo MRI of embryonic stem cells in a mouse model of myocardial infarction. Magn Reson Med 52, 1214-1219 (2004).
  95. Carr, C.A., et al. Bone marrow-derived stromal cells home to and remain in the infarcted rat heart but fail to improve function: an in vivo cine-MRI study. Am J Physiol Heart Circ Physiol 295, H533-542 (2008).
  96. Li, Z., et al. Comparison of reporter gene and iron particle labeling for tracking fate of human embryonic stem cells and differentiated endothelial cells in living subjects. Stem Cells 26, 864-873 (2008).
  97. de Vries, I.J., et al. Magnetic resonance tracking of dendritic cells in melanoma patients for monitoring of cellular therapy. Nat Biotechnol 23, 1407-1413 (2005).
  98. Zhu, J., Zhou, L. & Xing Wu, F. Tracking neural stem cells in patients with brain trauma. N Engl J Med 355, 2376-2378 (2006).
  99. Richards, J.M., et al. In vivo mononuclear cell tracking using superparamagnetic particles of iron oxide: feasibility and safety in humans. Circ Cardiovasc Imaging 5, 509-517 (2012).
  100. Gilad, A.A., et al. MR tracking of transplanted cells with “positive contrast” using manganese oxide nanoparticles. Magn Reson Med 60, 1-7 (2008).
  101. Gilad, A.A., et al. Artificial reporter gene providing MRI contrast based on proton exchange. Nat Biotechnol 25, 217-219 (2007).
  102. Cohen, B., Dafni, H., Meir, G., Harmelin, A. & Neeman, M. Ferritin as an endogenous MRI reporter for noninvasive imaging of gene expression in C6 glioma tumors. Neoplasia 7, 109-117 (2005).
  103. Zurkiya, O., Chan, A.W. & Hu, X. MagA is sufficient for producing magnetic nanoparticles in mammalian cells, making it an MRI reporter. Magn Reson Med 59, 1225-1231 (2008). Image reprinted with permission from John Wiley and Sons.
  104. Partlow, K.C., et al. 19F magnetic resonance imaging for stem/progenitor cell tracking with multiple unique perfluorocarbon nanobeacons. FASEB J 21, 1647-1654 (2007).
  105. Flogel, U., et al. In vivo monitoring of inflammation after cardiac and cerebral ischemia by fluorine magnetic resonance imaging. Circulation 118, 140-148 (2008).
  106. Flacke, S., et al. Novel MRI contrast agent for molecular imaging of fibrin: implications for detecting vulnerable plaques. Circulation 104, 1280-1285 (2001).
  107. Padhani, A.R. Dynamic contrast-enhanced MRI in clinical oncology: current status and future directions. J Magn Reson Imaging 16, 407-422 (2002).
  108. Caravan, P., et al. Collagen-targeted MRI contrast agent for molecular imaging of fibrosis. Angew Chem Int Ed Engl 46, 8171-8173 (2007).
  109. Mulder, W.J., et al. Molecular imaging of tumor angiogenesis using alphavbeta3-integrin targeted multimodal quantum dots. Angiogenesis 12, 17-24 (2009).
  110. Amirbekian, V., et al. Detecting and assessing macrophages in vivo to evaluate atherosclerosis noninvasively using molecular MRI. Proc Natl Acad Sci U S A 104, 961-966 (2007).
  111. Nahrendorf, M., et al. Activatable magnetic resonance imaging agent reports myeloperoxidase activity in healing infarcts and noninvasively detects the antiinflammatory effects of atorvastatin on ischemia-reperfusion injury. Circulation 117, 1153-1160 (2008).
  112. Huh, Y.M., et al. In vivo magnetic resonance detection of cancer by using multifunctional magnetic nanocrystals. J Am Chem Soc 127, 12387-12391 (2005).
  113. McAteer, M.A., et al. In vivo magnetic resonance imaging of acute brain inflammation using microparticles of iron oxide. Nat Med 13, 1253-1258 (2007).
  114. Artemov, D., Mori, N., Okollie, B. & Bhujwalla, Z.M. MR molecular imaging of the Her-2/neu receptor in breast cancer cells using targeted iron oxide nanoparticles. Magn Reson Med 49, 403-408 (2003).
  115. Yang, L., et al. Single chain epidermal growth factor receptor antibody conjugated nanoparticles for in vivo tumor targeting and imaging. Small 5, 235-243 (2009).
  116. Kelly, K.A., et al. Detection of vascular adhesion molecule-1 expression using a novel multimodal nanoparticle. Circ Res 96, 327-336 (2005).
  117. Johansson, L.O., et al. A targeted contrast agent for magnetic resonance imaging of thrombus: implications of spatial resolution. J Magn Reson Imaging 13, 615-618 (2001).
  118. Sosnovik, D.E., et al. Magnetic resonance imaging of cardiomyocyte apoptosis with a novel magneto-optical nanoparticle. Magn Reson Med 54, 718-724 (2005). Image reprinted with permission from John Wiley and Sons.
  119. Szczepaniak, L.S., et al. Myocardial triglycerides and systolic function in humans: in vivo evaluation by localized proton spectroscopy and cardiac imaging. Magn Reson Med 49, 417-423 (2003).
  120. Wolfe, C.L., et al. Assessment of myocardial salvage after ischemia and reperfusion using magnetic resonance imaging and spectroscopy. Circulation 80, 969-982 (1989).
  121. Jung, W.I., et al. 31P NMR spectroscopy detects metabolic abnormalities in asymptomatic patients with hypertrophic cardiomyopathy. Circulation 97, 2536-2542 (1998).
  122. Golman, K., et al. Cardiac metabolism measured noninvasively by hyperpolarized 13C MRI. Magn Reson Med 59, 1005-1013 (2008).
  123. Day, S.E., et al. Detecting tumor response to treatment using hyperpolarized 13C magnetic resonance imaging and spectroscopy. Nat Med 13, 1382-1387 (2007).
  124. Pruessmann, K.P., Weiger, M., Scheidegger, M.B. & Boesiger, P. SENSE: sensitivity encoding for fast MRI. Magn Reson Med 42, 952-962 (1999).
  125. Lustig, M., Donoho, D. & Pauly, J.M. Sparse MRI: The application of compressed sensing for rapid MR imaging. Magn Reson Med 58, 1182-1195 (2007).


Hide picture