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Molecular imaging Towards in vivo nuclear microscopy: iodine-125 imaging in mice using micro-pinholes Freek J. Beekman1, 2, David P. McElroy3, Frank Berger2, 3, Sanjiv S. Gambhir2, 3, Edward J. Hoffman3, Simon R. Cherry2, 4 1 Image

Sciences Institute, University Hospital Utrecht, E 02.222, Heidelberglaan 100, 3584 CX, The Netherlands Institute for Molecular Imaging, UCLA School of Medicine, Los Angeles, USA 3 Department of Molecular and Medical Pharmacology, UCLA School of Medicine, Los Angeles, USA 4 Department of Biomedical Engineering, University of California, Davis, USA 2 Crump

Published online: 19 April 2002 © Springer-Verlag 2002

Abstract. Position-sensitive gamma-radiation detectors equipped with collimators have been used for in vivo imaging of the distribution of radiolabelled molecules in laboratory animals and humans for several decades. To date, the best image resolution achieved in a rodent is on the order of 1 mm. Here we demonstrate how a basic and compact gamma camera can be constructed for in vivo radionuclide imaging in small animals, at much higher spatial resolution. Resolution improvements were obtained by combining dense, shaped, micro-pinhole apertures with iodine-125, an isotope with low energy emissions, ease of incorporation into a wide range of molecules, and straightforward translation into the clinic via other isotopes of iodine that are suitable for nuclear medicine imaging. 125I images of test distributions and a mouse thyroid have been obtained at a resolution of as high as 200 µm using this simple bench-top camera. Possible future applications and extension to ultra-high-resolution emission tomography are discussed. Keywords: Iodine – Pinhole imaging – SPET – Thyroid – Mouse Eur J Nucl Med (2002) 29:933–938 DOI 10.1007/s00259-002-0805-6

Introduction The increasing availability of genetically modified mice enables a wide range of human diseases to be studied in Freek J. Beekman (✉) Image Sciences Institute, University Hospital Utrecht, E 02.222, Heidelberglaan 100, 3584 CX, The Netherlands e-mail: [email protected] Tel.: +31-30-2507779, Fax: +31-30-2542531

European Journal of Nuclear Medicine Vol. 29, No. 7, July 2002

animal models. Radiolabelling of small molecules, antibodies, peptides and probes for gene expression, in concert with dedicated imaging equipment, enables molecular imaging in vivo. Until recently, most studies of this type required animal dissection and specimen counting, animal sectioning followed by autoradiography, or other assays that involve dissection or sectioning and counting. Groups of animals are sacrificed at a range of selected time points after radiolabelled tracer administration to determine the uptake and retention of the tracer in the body. Such approaches preclude longitudinal studies in the same animal and the use of individual animals serving as their own control during interventional studies. Imaging of radiolabelled agents in living intact animals or humans can be performed using positron emission tomography (PET) [1] with positron-emitting radionuclides, or using a gamma camera equipped with a collimator for the study of single gamma-emitting radionuclide distributions [2]. Gamma camera imaging can also be performed in tomographic mode by acquiring projection images of the subject at different angles, followed by tomographic image reconstruction. The latter is most often referred to as single-photon emission tomography (SPET) [3]. Up to now, dedicated small animal PET and gamma camera imaging assays have been limited to a spatial image resolution of approximately 2 mm [4] and 1 mm [5], respectively. Pinhole gamma camera imaging (Fig. 1) offers the ability to obtain high-resolution images of gamma-emitting tracers in small objects like the human thyroid and small animals [6, 7, 8]. However, a fundamental factor limiting image resolution is the penetration of gammaradiation along the edge of the pinhole collimator. Gamma-rays and X-rays with low energies such as those emitted by iodine-125 (27–35 keV) are less penetrating than those emitted by technetium-99m (140 keV) or iodine-123 (159 keV). Therefore, smaller diameter pinholes can be used for such low photon energies, poten-

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tially increasing the image resolution. To further reduce radiation penetration of the aperture, material with a higher stopping power than common materials such as lead or tungsten can be used (Fig. 2). Note that gammaand X-rays emitted by 125I have an average path-length of about 1 cm in soft tissue, which allows a sufficient fraction of emitted radiation flux to escape from the mouse for external imaging. For most mouse imaging situations, this detectable fraction is higher than that of 99mTc or 123I gamma rays in human imaging. 125I is a readily available radionuclide with a half-life of 60.2 days. 125I-labelled tracers are commonly used in molecular biology, and many 125I-labelled nucleic acids, antibodies, ligands, and other radiopharmaceuticals [e.g. 9, 10] are commercially available. In addition, many other 125I-labelled imaging probes can be readily prepared using commercial iodination reagents and kits. The availability of 125I-labelled tracers has motivated several groups to develop dedicated systems capable of in vivo imaging of 125I in small animals [e.g. 11, 12]. An important advantage of iodine imaging is that iodine-labelled molecules provide a convenient bridge from animal models up into the human. One can replace 125I by 123I (half-life 13.2 h), whose 159-keV gammarays have a sufficiently long path-length for imaging in humans. This substitution does not change the chemical properties of the imaging probe. Furthermore, imaging of iodine-labelled molecules by PET can be accomplished with the positron emitter 124I. There is also the prospect of using iodine for radionuclide therapy, where the 125I used in the imaging probe is replaced by 131I, which emits both beta- and gamma-rays. A wide range of 123I- and 131I-labelled molecules are currently available for use in humans [13], and there are several recent papers indicating that iodine-labelled agents will play an even more important role in biomedical research, diagnosis, and (radionuclide) therapy in the future. Examples include the early diagnosis of Alzheimer’s disease by targeting the amyloid plaque [14], and the NaI symporter gene, which has potential as a PET and SPET reporter gene or even for therapeutic approaches [15, 16]. A high-resolution gamma camera for in vivo imaging of 125I-labelled biomolecules and SPET systems based on such a camera therefore have the potential to be important new tools in the development of new imaging probes, and in testing new therapeutic approaches in mouse models of human disease. Because of the translational ability afforded by the isotopes of iodine discussed above, diagnostic and therapeutic approaches can be moved from mouse to man with relative ease.

Materials and methods Pinhole gamma-camera imaging. Pinhole collimation was one of the first collimation methods ever used for gamma-ray imaging [17]. Pinhole collimation can offer relatively high quality images in small

organs and small animals compared with the standard parallel hole collimation typically employed in clinical studies. The characteristics of pinhole imaging are discussed below with the aid of mathematical equations (e.g. from [18]), for system dimensions which are representative for the present ultra-high-resolution camera. The decrease in quality of pinhole images for larger organs or organs deeper in the body is partly explained by the decrease in count sensitivity S with increasing distance z (Fig. 1) between the emitting source and pinhole; For a point source located in the centre of the field of view S≈D2/16z2, where D is the hole diameter. For objects close to the collimator, the efficiency of a pinhole collimated system can be orders of magnitude higher than that of a parallel hole system. The spatial resolution decreases with increasing pinhole to source distance, particularly when a detector with relatively low intrinsic resolution is used. The resolution of an ideal pinhole system (assuming no transmission of gamma- and X-rays by the aperture material, and assuming a detector with an ideal resolution) can be approximated by the geometric resolution Rg: (1) where l is the distance between the pinhole and the detector. For an l value of 100 mm, the resolution would decrease by only about 5% when increasing z from 5 to 10 mm. Unfortunately, the intrinsic resolution Ri of the detector can reduce the total resolution Rt dramatically, in particular when a larger object to pinhole distance or a low magnification factor (l/z) is used. On assuming Gaussianshaped kernels for intrinsic resolution and pinhole blurring, the expression for Rt in terms of the full-width at half-maximum (FWHM) [19] reads: (2) This equation shows that increasing the magnification factor reduces the loss of image resolution due to intrinsic detector blurring. The magnification factor, however, is often limited by the detector surface area available. The resolution for a 100-µm pinhole reduces from 105 µm for an ideal detector to 183 µm when Ri is 3 mm, z=5 mm and l=100 mm. At an increased z value of 10 mm, the resolution degradation due to detector blurring is more dramatic: instead of 110 µm for an ideal detector, a resolution as bad as 319 µm is obtained. The width W of the field of view of a pinhole system with a sufficiently large detector behind it is determined by the opening angle α (Fig. 2a) of the pinhole system and the object distance to the pinhole z: (3) For the 30º pinhole opening angle used in the present system we obtain a width of the field of view which equals the distance between the object and the pinhole opening. Because of the excellent imaging characteristics close to the pinhole, pinhole systems have been mainly used for imaging the human thyroid and small animals [e.g. 5, 6, 7, 8,20]. From Eq. 1, we see that an image resolution approaching the pinhole diameter can be obtained when the pinhole to detector distance is sufficiently large and the object is close to the pinhole. The smallest pinhole diameters reported so far are about 0.5 mm, and have resulted in a SPET image resolution of about 1 mm [6, 7, 8]. Pinhole apertures. A schematic of the pinhole collimated camera is shown in Fig. 1, and details of a pinhole aperture made from a


935 Fig. 3. Desktop pinhole camera with mouse during acquisition of a thyroid image

Fig. 1. Schematic of camera system with pinhole aperture. The radiation emitted by an organ or tumour enters the camera through the pinhole. The image of the organ is enlarged on the detector, which facilitates imaging at a spatial resolution higher than the detector resolution

gold-platinum alloy, as used in the present system, are shown in Fig. 2. Fig. 2a shows a cross-section through the design of the pinhole apertures. The apertures were made from a 90% Au–10% Pt alloy (diameters 100 µm and 250 µm, thickness 1.5 mm, α 30º). The pinholes were manufactured using electrical discharge machining (EDM) by Optimation, Midvale (UT). EDM is based on the use of high-voltage pulses between the workpiece (pinhole) and the electrode (“drill”), which are both in a dielectric fluid. At each pulse a thin layer of material is removed. Each pinhole has a straight channel which is approximately 0.6 times the diameter (Fig. 2a). It has been shown that pinhole designs with such a channel may avoid excessive radiation leakage along the pinhole edge as compared to pinholes with a knife-edge design [21]. Fig. 2b shows a scanning electron microscope image of the pinhole aperture with a 100 µm diameter and 60 µm channel length. Gamma detector. In standard clinical gamma cameras, scintillation crystals made of NaI(Tl) with thicknesses of the order of 10 mm are commonly used as a scintillator. These are covered by an aluminium layer of about 0.75 mm to avoid light exposure and hydration of the crystal. In the case of the low energy radiation of 125I, a layer of this thickness would absorb or scatter a significant fraction of the gamma-rays, while the large crystal thickness (required to stop a sufficiently high fraction of gamma-rays emitted from the clinically useful, higher energy isotopes such as 99mTc or 123I) would unnecessarily decrease the intrinsic resolution for 125I. In order to tackle these problems, a thin (1.5 mm) NaI(Tl) crystal, covered by a 0.25-mm-thick layer of aluminium was used. A 77×77 mm position-sensitive photon multiplier tube (Hamamatsu, type R2487-02, with crossed wire anode construction) was coupled to the crystal for position determination. The average intrinsic detector resolution on the camera surface was measured to be 3.1 mm. The entire imager was shielded using a box of 5.5-mmthick lead walls, with outer box dimensions of 91×91×273 mm (Fig. 3) The distance l between the pinhole and the detector is 105 mm.

Fig. 2. a Cross-sectional drawing through the gold-platinum alloy pinhole aperture with a diameter D. b Scanning electron microscope image of a pinhole aperture (100 µm diameter with 60 µm channel length) constructed by electron discharge machining


Assessment of image resolution. For measuring image resolution, two line sources were used (specific activity 0.3 µCi/mm), consisting of rectangular (40×200 µm) borosilicate capillary tubes (Friedrich and Dimmock, Philadelphia, Pa.), and filled with NaI (125I) using capillary action. The two capillaries were separated by 1.05 mm using a plastic spacer. The short sides (40 µm) of the capillaries were directed towards the pinhole during image acqui-

936 Fig 4. Details of images and associated counting density profiles of two 40-µm-wide 125I capillaries separated by 1.05 mm, produced by a 250µm pinhole (a) and a 100-µm pinhole (b). The distance between the capillary tubes and the pinhole opening was 5 mm

sition. The FWHM was calculated from count density profiles (Fig. 4) drawn through the image of the sources. Mouse imaging. Animal studies were performed with protocols approved by the UCLA Animal Research Committee. 125I thyroid images of a 25-g Swiss Webster mouse were acquired 4 h after tail vein injection of 0.25 mCi NaI (125I). In order to avoid unnecessary attenuation of the emitted photon flux, the lucite cylinder (Fig. 3) has a hole in the vicinity of the thyroid. The mouse was anaesthetised by i.p. injection of ≈40 µl of a ketamine and xylazine (4:1) solution, just before imaging. Fig. 5a was obtained with the 250-µm pinhole (20 min image acquisition time) and Fig. 5b was obtained with a 100-µm pinhole (50 min image acquisition time). A histological slide (Fig. 5c) was prepared to compare the shape of the thyroid in the 125I images with anatomy. In order to prepare the histological slide, the trachea was fixed in 10% formaldehyde and dehydrated in upgrading alcohols and xylene before going in paraffin for cutting of sections for the glass slide. Tissue on the slide was stained with haematoxylin and eosin.

Results Two micro-pinholes with apertures of 100 µm and 250 µm were successfully fabricated using electrical discharge machining (Fig. 2b). Line source images at a distance of 5 mm from the pinhole are shown for 100-µm and 250-µm pinhole diameters (Fig. 4). The measured image resolution in terms

of FWHM was 190 µm for the 100-µm pinhole and 288 µm for the 250-µm pinhole. After correcting for the additional blurring caused by the thickness of the line sources, these resolutions improved to 186 µm and 285 µm, respectively. The theoretical values of 185 µm (100-µm pinhole) and 303 µm (250-µm pinhole), calculated using Eq. 2, are in good agreement with the measurements. In vivo 125I thyroid images of a Swiss Webster mouse, approximately 4 h after tail vein injection of 0.25 mCi 125I, were acquired (Fig. 3) and are shown in Fig. 5a and b. The images correspond to an area of just 1.0×1.2 mm. Fig. 5a was obtained with the 250-µm pinhole (20 min image acquisition time) and Fig. 5b was obtained with a 100-µm pinhole (50 min image acquisition time). The similarity in shape of the thyroid in these 125I images with anatomy is illustrated with a histological slide of a slice showing the thyroid of a mouse (Fig. 5c). A striking detail is the cold indents on the left and right sides of the thyroid images, which likely correspond to the parathyroids. In contrast with the thyroid, the parathyroids do not have any significant iodine uptake [22]. The sizes of the parathyroids as estimated from histology are about 100×250 µm and 70×125 µm.


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Fig 5. 125I image of a mouse thyroid made with a 250-µm pinhole (a) and a 100-µm pinhole (b). The thyroid of a mouse has a size of about 1 square millimetre. c Histology: microscope image of a slice through the mouse thyroid. Note the agreement in shape with the 125I thyroid scan and the small areas with reduced 125I uptake on both sides of the thyroid, corresponding to the positions of the parathyroid glands (arrows)

Discussion The feasibility of ultra-high-resolution in vivo imaging of single photon-emitting isotope distributions was investigated. Combining tiny high-density pinholes with a low-energy radio-isotope (125I) allows details of a few hundreds of µm to be resolved in the thyroid of a living mouse, even when a very basic position-sensitive gamma detector is used. In this work, gold alloy pinholes were used instead of common lead or tungsten. Other groups have tried to de-


crease radiation leakage for higher energy photons like those of 131I or even of 511-keV photons by using depleted uranium [23, 24]. Uranium pinholes have an even higher stopping power than gold, but they are hard to manufacture and not commonly available. Furthermore, there are risks that toxic uranium oxide may be shedded, which requires the performance of wipe tests and plating. In addition, uranium is radioactive, which causes a radiation background on the detector that can be a significant problem [23]. According to Eq. 2, the spatial resolution of our imager may be improved by tens of percent by using a detector with a higher intrinsic resolution. To achieve such an improvement, segmented crystals could be used instead of the continuous crystal used in the current set-up. A further increase in intrinsic resolution is possible with small highresolution semi-conductor detectors, with a resolution in the order of a few hundred micrometres; such detectors are currently being developed by several universities and companies. According to Eq. 2 the higher intrinsic resolution will also strongly reduce resolution loss with distance. Since the count sensitivity of pinhole imagers is known to decrease quickly with increasing source distance, the most likely area of application of the device presented herein is in the imaging of small structures located close to the surface of the animal, such as the thyroid, the breast and small structures in the brain. A particularly important application could be the use of this technology to study tumour models, particularly where tumours are located just under the skin. At this resolution, much information could be obtained about regional tumour microvasculature and perfusion, delivery and retention of drugs and other molecules targeted to cancer cells. The ability to study these models repeatedly and non-invasively in an intact animal presents obvious opportunities. With the aim of extending our application to three-dimensional tomographic imaging (SPET), we anticipate combining high-resolution pinhole collimators with small high-resolution detectors such as cadmium-zinc-tellurium (CdZnTe) arrays [e.g. 25]. The small detector size which can be used with high-resolution detectors will facilitate surrounding a mouse with hundreds of pinholes, thus increasing the count sensitivity dramatically. Based on the image resolution obtained in this study and on resolution and sensitivity calculations, we believe such small detectors will allow tomographic in vivo imaging of 125I at a resolution of a few hundred micrometres. Systems with many pinholes have been proposed previously for small animal imaging and have already resulted in a SPET image resolution of about a millimetre for 99mTc [20]. Using the technologies presented in this paper, in combination with the advantageous properties of 125I [suitability for pinhole imaging, flexibility of labelling a wide variety of imaging probes and extension into human imaging with 123I and 125I, and with 131I for therapy], offers many opportunities. The future development of ultra-high-resolution SPET systems will allow the

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three-dimensional assessment of the distribution of a wide range of radiolabelled tracers. Labour-intensive in vitro or ex vivo methods may be replaced in several cases by the imaging of intact living animals at the submillimetre level. This will also allow for accurate longitudinal study designs, and may accelerate the development of new diagnostic or therapeutic agents. Importantly, in vivo imaging with iodine leads to many translational opportunities. By utilising different iodine isotopes, non-invasive diagnostic imaging and therapeutic approaches developed in the mouse can be translated directly to the clinic, where conventional radionuclide imaging techniques such as PET and SPET and radio-iodine therapy are already in common use. Acknowledgements. The authors thank Aune T.M. More, Department of Human Genetics (UCLA), for providing the histology slide of the mouse thyroid, Ken Meadors (UCLA) for advice and for machining parts of the imager, Sander van Geloven (Utrecht University) for computer support, and Dean Jorgenson, Optimation, Midvale (UT), for manufacturing pinholes. We thank Dr. Magnus Dahlbom and Collin Dimmock (UCLA) for providing NaI(125I) and phantom preparation, and Dr. Yiping Shao and Dr. Tatsushi Toyokuni (UCLA) for helpful discussions. This work was supported in part by the Netherlands Organization for Scientific Research, Grant R90-172.

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