Small-animal molecular imaging methods

Abstract

The ability to trace or identify specific molecules within a specific anatomic location provides insight into metabolic pathways, tissue components, and tracing of solute transport mechanisms. With the increasing use of small animals for research, such imaging must have sufficiently high spatial resolution to allow anatomic localization as well as sufficient specificity and sensitivity to provide an accurate description of the molecular distribution and concentration.

Methods: Imaging methods based on electromagnetic radiation, such as PET, SPECT, MRI, and CT, are increasingly applicable because of recent advances in novel scanner hardware and image reconstruction software and the availability of novel molecules that have enhanced sensitivity in these methodologies.

Results: Small-animal PET has been advanced by the development of detector arrays that provide higher resolution and positron-emitting elements that allow new molecular tracers to be labeled. Micro-MRI has been improved in terms of spatial resolution and sensitivity through increased magnet field strength and the development of special-purpose coils and associated scan protocols. Of particular interest is the associated ability to image local mechanical function and solute transport processes, which can be directly related to the molecular information. This ability is further strengthened by the synergistic integration of PET with MRI. Micro-SPECT has been improved through the use of coded aperture imaging approaches as well as image reconstruction algorithms that can better deal with the photon-limited scan data. The limited spatial resolution can be partially overcome by integrating SPECT with CT. Micro-CT by itself provides exquisite spatial resolution of anatomy, but recent developments in high-spatial-resolution photon counting and spectrally sensitive imaging arrays, combined with x-ray optical devices, hold promise for actual molecular identification by virtue of the chemical bond lengths of molecules, especially biopolymers.

Conclusion: Given the increasing use of small animals for evaluating new clinical imaging techniques and providing more insight into pathophysiologic phenomena as well as the availability of improved detection systems, scanning protocols, and associated software, the sensitivity and specificity of molecular imaging are increasing.

FIGURE 1

Micro-PET images of [18F] FDG distribution in a 25g mouse showing whole-body (A), cardiac ECG-gated (B) and cardiac dynamic (C) distributions obtained simultaneously with list-mode imaging. Whole-body and cardiac dynamic data from the Inveon DPET. ECG-gated images from the LabPET courtesy of R. Lecomte, Université de Sherbrooke.

FIGURE 2

Evolution of micro-PET scanners for mouse imaging. Sensitivity increases with solid angle coverage and detector depth (*value based on simulation). Focus 220 and Inveon DPET figures from Bao et al, 2008 with permission. Tapered detector design courtesy of S.St.James and S.Cherry, UC Davis.

FIGURE 3

Detector point spread function (PSF) and positron range model (β+) can be included in the system matrix used for iterative image reconstruction (A), modified from Shaart et al, 2008 with permission. Improved image resolution (B) is obtained compared to filtered backprojection (FBP) using maximum a posteriori (MAP) reconstruction with PSF and positron range modeling (MAPR). With MAPR the renal cortex can be resolved clearly in the mouse kidney (C), with improved accuracy validated against ex-vivo biodistribution (BioD) values. 61Cu-PTSM mouse images from the micro-PET Focus courtesy of R. Laforest, Washington University.

FIGURE 4

Parametric images of regional myocardial glucose uptake (rMGU) influx rate constant (Ki) in the mouse after PSF reconstruction (A), simulated image from Spinelli et al, 2008a with permission. An input function and temporal basis functions (tissue response) for fully 4D image reconstruction can be derived directly from the PET raw data (B), from Reader et al, 2007 with permission. Recovery of temporal resolution (early 10s frame) with 4D vs. conventional 3D reconstruction of FDG brain images from the HRRT (C), images courtesy of A.J. Reader, Montreal Neurological Institute.

FIGURE 5

End-diastolic and end-systolic mid-ventricular short-axis MR images of the mouse heart, taken from a cine data set with 16 frames across the cardiac cycle. The spatial resolution was 0.15 × 0.15 × 0.5 mm³, and the temporal resolution was 6 ms. The images were acquired using a 7T MR system, and the total scan time was approximately 5 minutes.

FIGURE 6

End systolic short-axis cine DENSE displacement (A,B) and circumferential shortening (C,D) maps of the normal (A,C) and 7-day postinfarct (B,D) mouse heart. The region of infarction (11:00–5:00) has markedly reduced displacement (B) and shortening (yellow-orange) (D).

FIGURE 7

Gadolinium-enhanced inversion-recovery T1-weighted MR image of the mouse heart 1 day after induction of experimental myocardial infarction. The hyperintense region delineates the infarcted tissue (arrows). These images can be used to define the infarct, adjacent, and remote zones. Also, multislice data sets that cover the entire left ventricle can accurately assess infarct size.

FIGURE 8

(A) Example 3D MRA maximum intensity projection of the aortic arch. The yellow line represents the measurement plane for phase contrast imaging, the red arrow shows the inner radius of the aortic arch, and the blue arrow shows the outer radius. (B) Wall shear stress as a function of cardiac phase for the inner and outer radii for a group of 5 mice.

FIGURE 9

Different approaches to combined PET/MRI. (A) Long optical fibers used to couple scintillators to PMTs residing outside the magnet. (B) Avalanche photodiodes (APDs) coupled directly to scintillator elements. (C) Scintillators coupled through short optical fibers to APDs. In all cases, the scintillators are centered axially in the MRI field of view.

FIGURE 10

(A) Fibrin targeted contrast agnet (EP-2104R) consisting of an 11 amino acid peptide functionalized with 2 GdDOTA-like moieties at both the C-and the N-terminus of the peptide. (B) Bright spot thrombus MR imaging with the fibrin-targeted contrast agent in a rat stroke model. The clot in the internal carotid and middle cerebral arteries is demonstrated (arrows) (Images courtesy of Peter Carvavan, Martinos Center, MGH).

FIGURE 11

Sequentially acquired PET and MR data in a myocardial infarction rat model. The decreased FDG uptake area in the anterior wall correlates with the hyperenhancing region demonstrated by the contrast-enhanced MRI (ceMRI) study (arrows). The short-axis images from apex (left) to base (Right) as shown. (Reproduced with permission (50))

FIGURE 12

Comparison of the imaging geometries of a conventional parallel-hole collimator and a pinhole collimator in preclinical imaging of small animals. Through magnification of the small object at close distance onto a larger detector area, the pinhole collimation geometry offer substantial increase in detection efficiency as compare to the parallel-hole collimation geometry and the spatial resolution is determined by the size of the pinhole aperture.

FIGURE 13

Comparison of the spatial resolution of a typical low-energy high-resolution (LEHR) parallel-hole collimator with three pinhole collimators with different pinhole aperture size as a function of source distance. Comparison of the geometry efficiency of the same collimators as a function of source distance. The graphs in (left panel) and (right panel) show the advantages of pinhole collimation over parallel-hole collimation when imaging small animals at close distance.

FIGURE 14

A multi-pinhole collimation geometry where the projections of a common field-of-view (indicated by the circle defined by the intersects of the three projections) share the same detector area without overlap. The increase number of projections provides increase geometric detection efficiency. However, the geometry detection efficiency from each pinhole aperture is less that of the single pinhole collimation where the entire detector area can be used with large magnification.

FIGURE 15

Transaxial images at end-diastole from a gated myocardial perfusion micro-SPECT study of a rat. The animal was injected with ~5 mCi of Tc-99m labeled Sestamibi and imaged using a dual detector Gamma Medica-Ideas XSPECT system with each detector fitted with a single pinhole collimator with 1 mm diameter aperture size. The total acquisition time is 20 minutes and the reconstructed images were post-processed using a Butterworth filter with a cut-off frequency of 0.1 cycle/voxel and order 8. The images indicate the ability of the micro-SPECT system to clearly delineate the myocardium showing normal perfusion.

FIGURE 16

(a)From left to right, sample raw transaxial image from the gated myocardial perfusion micro-SPECT study of a rat in Figure 15 using a 3-D pinhole filtered backprojection image reconstruction method, and the same transaxial image post-processed with a Butterworth filter with order 8 and cut-off frequencies of 0.15 and 0.1 cycle/voxel.(b) from left to right, the reconstructed image from the same transaxial slice as in (a) obtained from using an iterative 3-D OS-EM pinhole image reconstruction method without any correction at 2, 5, 8 and 11 iterations, and post-processed using a Butterworth filter with order 8 and cut-off frequencies of 0.1 cycle/voxel.(c)from left to right, the reconstructed image from the same transaxial slice as in (a) obtained from using an iterative 3-D OS-EM pinhole image reconstruction method with correction of the geometric response of the pinhole collimator at 2, 5, 8 and 11 iterations. The images demonstrate the effectiveness of the corrective pinhole image reconstruction methods in improving the reconstruction image quality by improving image resolution without noise amplification. Reproduced with permission (62))

FIGURE 17

Sample ultra-high resolution ^99mTc-tetrofosmin SPECT images of a mouse's heart in end-diastole (ED) (left) and end-systole (ES) showing myocardial perfusion in tiny details such as the papillary muscles and the right ventricular wall. The 30g male C57Bl/6 mouse was injected intravenously with 190 MBq of ^99mTc-tetrofosmin and anesthetized using ketamine-medetomidine-atropine. At 45 min after injection, the mouse was imaged for 1 hour using the U-SPECT-II system with 0.6mm diameter pinhole inserts. During image acquisition, an electrocardiogram trigger signal was acquired (BioVet; m2m imaging) and incorporated in the list-mode data.A 16-gate reconstruction was performed. Image data courtesy of Freek J. Beekman, Ph.D. (Reproduced with permission (69))

FIGURE 18

Volume rendered display of a 3D micro-CT image of the coronary arteries of a mouse heart wall. A single arterial tree is highlighted in red. (Courtesy of G.S.Kassab)

FIGURE 19

Volume rendered display of the contrast-billed microvasculature in a mouse ear. Upper panels show vessels at Day 4 after local injection of VEGG. The read vessels are all vessels less than 60μm diameter. Lower panels show a mouse ear 10 days after local injection of VEGF. Note, the great increase in the highlighted small vessels. (Courtesy of Dr.J.A. Nagy)

FIGURE 20

Left panel – micro-CT image of a transverse section through a mouse's lower thorax and upper abdomen. Contrast fills the major blood vessels and the stomach. The blood vessels contain iodine and the stomach contains barium-based contrast agent. Multi-energy x-ray allows discrimination of the iodine, barium and normal tissues (green). Right panel is color-coded to show the location of the barium and iodine contrast agents. (Reproduce with permission (87))

FIGURE 21

Left panel – CT image of test phantom used to evaluate coherent x-ray scatter image contrast of various polymers. Mid panel – same tomographic slice generated from coherent scatter of a 17.5keV x-ray beam at 12.2° from the illuminating beam. Right panel – same tomographic slice generated from coherent scatter at 22keV and 7.3°. Note, the removal of contrast of the polymers with illuminating energy. The coherent scatter images are blurred in the top-to-bottom direction because the phantom was scanned at 300μm increments in that direction, whereas the regular CT scan is isotropic at 20μm voxel resolution.