S. J. Riederer

Selective iodine imaging using K-edge energies in computerized X-ray tomography

Iodine is commonly used as a contrast material in computerized x-ray tomography. In some cases the determination of the iodine distribution in the image may be prevented by the presence of bone or tissue variations within the tomographic slice. This paper describes a method for quantitative selective imaging of the iodine concentration in the slice. The method employs scans using three heavily filtered x-ray beams, two having mean energies which straddle the iodine K edge (33 keV) and another at a slightly higher energy. The results are independent of tissue and bone over a broad range of projection path lengths. It is shown that, for separation of iodine from one other material, a two-beam K-edge approach requires less integral dose than a two-beam technique at conventional CT energies for slice diameters up to 30 cm. For selective iodine imaging in the presence of more than one other material, the three-spectrum K-edge technique is a necessity. Exposure requirements and beam-hardening corrections are discussed in detail and a computer-simulated CT image generated by the proposed scheme is presented.

Relative properties of tomography, K-edge imaging, and K-edge tomography.

The properties of tomography, K-edge imaging, and K-edge tomography are discussed in relation to the imaging of small concentrations of elements such as iodine and xenon and are compared by means of phantom images. It is demonstrated that the complementary selectivities provided by depth and energy subtraction are combined in K-edge tomography. Using a three-spectrum subtraction technique, the iodine difference signal predicted by computer calculations is on the order of 8000 times that of an equal concentration fo bone. The corresponding ratio in tomography without energy subtraction is 20:1. It is argued that K-edge tomography can successfully eliminate artifacts due to tissue inhomogeneities which presently enable 0.6% variations in tissue attenuation to mimic minimum detectable iodine signals in conventional computed tomography. Various instrumentation possibilities and energy subtraction techniques are discussed.

Spectral considerations for absorption-edge fluoroscopy.

In our previous reports on absorption-edge fluoroscopy, it was not possible to relate fully the subtleties involved in the selection of spectral parameters. This paper is intended as an overview of this important aspect of the technique. It is shown that, by using the 1-kVp, 2-filter technique, it is possible to image certain elements (e.g., iodine and xenon) in the presence of tissue variations of +/-2 cm about the thickness at which perfect tissue cancellation takes place. Use of logarithmic signal processing extends this range, but bone thickness variations may not be accomodated because only two x-ray energies are involved in the imaging process. Use of a 3-kVp, 3-filter technique with logarithmic signal processing is shown to solve this problem. Computer simulations show that 1-mg/cm2 iodine may be imaged in the presence of 10 cm or more tissue variations and 2000-mg/cm2 bone variations.

Limitations to iodine isolation using a dual beam non-K-edge approach.

In dual-beam selective iodine imaging, images of an object are made with each of two spectrally different x-ray beams. The mean beam energies may either straddle the 33 keV iodine K-edge or both lie above the K-edge. Both patient exposure considerations and the availability of sufficient x-ray flux make the latter approach favorable for tissue thicknesses exceeding 5 cm. Consider such an approach in which image contrast from tissue is suppressed in the difference image. It is proven theoretically that the residual bone-to-iodine contrast is a constant independent of the two mean beam energies used. This invariance principle is demonstrated experimentally by comparing images made from different pairs of x-ray spectra. Observed contrast ratios match the predicted value very well. In dual-beam imaging, contrast from only one material may be suppressed. Other substances yield residual signals which compete with the iodine. Subtleties of this incomplete cancellation are demonstrated, discussed, and quantitated. A contrast enhancement factor (CEF) is defined as the factor by which iodine contrast is enhanced in a multiple beam subtraction technique relative to monoenergetic imaging at 40 keV. CEFs are determined for tissue and bone cancellation separately and their limits are discussed. Images of a simulated artery containing iodine superimposed over a Rando head and neck phantom show that the CEF limitation for dual beam imaging is quite severe compared to a time dependent mask mode imaging approach. Finally, optimum, energies for dual beam images are discussed.

Three-beam K-edge imaging of iodine using differences between fluoroscopic video images: Experimental results

In an earlier article we discussed the rationale for using differences between video images in three-beam selective iodine K-edge imaging. Rather than combining three initial images Li linearly to yield the final image k1L1 + k2L2 + k3L3, differences between the Li were first generated and then combined either to linear or quadratic order. This approach was motivated by the desire to suppress the large multiplicative biases of fluoroscopic imaging and justified by theoretically proving that k1 + k2 + k3 is approximately equal to 0. In this paper we discuss the instrumentation and experimental results obtained from this difference-based technique. A specially-constructed apparatus is described which automatically selects the optimum combination coefficients and combines the difference images up to quadratic order at realtime video rates. Three methods for generating K-edge subtraction images are compared: the former approach in which the Li are linearly combined and combination of differences to linear and quadratic order. In imaging phantoms in which the iodine distribution is known, the resultant subtraction images from all three methods appear similar. Inspection of signal sizes shows that the quadratic difference-based approach provides superior bone and tissue residual suppression by about a factor of 2. In imaging phantoms in which the iodine distribution is unknown, incomplete suppression of x-ray scatter and image intensifier veiling glare prevent a quantitative comparison of performance of the three algorithms. An experiment verification is provided of the theorem which states that k1 + k2 + k3 is approximately equal to 0.

Three‐beam K‐edge imaging of iodine using differences between fluoroscopic video images: Theoretical considerations

Our lab has previously generated selective iodine images with an image intensifier fluoroscopic system using a three-beam K-edge approach. Logarithmically amplified video images Li were linearly combined to yield the final image k1L1 + k2L2 + k3L3. This paper discusses refinements of the K-edge technique. A study is made of the manner in which contrast-reducing effects such as x-ray scatter and image intensifier veiling glare enter into the final image. If such biases can be approximated as multiplicative and independent of the x-ray spectrum, and if the sum of the ki is zero, then the biases are canceled. Experimental data is presented which demonstrates that the inaccuracy due to such biases can be reduced by a factor as large as 10. The theorem that K1 + K2 + K3 is approximately equal to 0 is proven rigorously and discussed. Because the ki add to zero, the final image can be expressed as a linear combination of two of the differences between the Li. A difference-based algorithm which reduces biases and make allowance for nonlinearities such as beam hardening is proposed and discussed.

Application of the noise power spectrum to positron emission CT self-absorption correction.

Two methods are compared in correcting for self-absorption in positron emission CT scans, or equivalently, in determining the integral of attenuation along a strip in a cross section. These are CT reconstruction and direct measurement. It is shown that the former method is slightly more precise than the latter with the degree of improvement proportional to the number of angles used. Additionally, it is demonstrated that the CT approach is statistically very inefficient in its use of detected events in comparison to the direct approach. For the CT case, it is shown using the noise power spectrum that two-thirds of the variance in the attenuation integral is from one projection alone, namely that along the direction of the strip. The remaining one-third is primarily from adjacent projections. For the problem of determining the attenuation in a single strip within the slice, the CT approach is not recommended. However, if the problem is extended to a complete set of strips within the slice, as is desired in positron emission CT, the CT approach may require considerably less counting time than the direct approach for comparable precision.

Real-Time Computerized Fluoroscopic Cardiac Imaging

A computerized fluoroscopy system employing several time and energy subtraction algorithms has permitted good visualization of the cardiovascular system using peripheral intravenous iodine injections of about 1 cm3/kg. Image contrast im-provements of 8-16 over conventional fluoroscopy are common. Several canine and human imaging studies are described including visualization of myocardial infarc-tions as regions of anomalous image grey shade. The system employs a standard image intensified fluoroscopy system and a specially constructed real-time image processor. Quasimonoenergetic x-ray beams formed by filtration deliver typical doses of 400 mR/sec in adult human cardiac exams.

SYSTEM FOR AUTOMATED REAL-TIME GENERATION OF HIGHER ORDER ENERGY SUBTRACTION IMAGES IN DIGITAL FLUOROSCOPY.

When energy subtraction methods are employed in digital radiography it is necessary to combine the initial images in a weighted fashion. This paper discusses the instrumentation developed to automatically generate up to quadratic order combinations of x-ray images at realtime video rates in energy subtraction digital fluorography. Although the energy mode used is three-beam K-edge subtraction, the same instrumentation can also be used for dual beam imaging. The method is implemented by first acquiring logarithmically amplified digital fluorographic images of a calibration phantom with each of three spectrally different x-ray beams. The distribution of iodinated contrast material in the phantom is presumed known and the phantom thicknesses cover the range expected to be encountered clinically. Next, individual pixel values from the images are automatically fed to a PDP 11/10 computer where the weighting coefficients are calculated by least squares. The optimum coefficients are next loaded from the computer into a hardware multiplier network capable of combination of 256 x 256 pixel images at realtime video rates. Once loaded with the fitted coefficients the network is then calibrated and ready to generate energy sub-traction images of arbitrary objects. The complexity of the image combination may range from linear to full quadratic combination of differences of the three initial images. The emphasis of this paper is a discussion of the closed form iodine specific fit used to determine the coefficients and the instrumentation aspects of the system: design of the computer interface and design and operation of the multiplier network. Sample experimental results are presented.

Real-Time Computerized Fluoroscopy and Radiography—A Progress Report

This report describes our initial experiences in the medical profession with real-time digital processing of X-ray transmission information taken from the video output of a conventional X-ray fluoroscopy and cine fluoroscopy system. X-rays are produced with an X-ray tube operated between 45 and 100 kVp, 1 to 300 mA, and filtered by a variety of materials, the choice depending on the type of information which is to be isolated during the imaging process. The transmitted beam is detected by a cesium iodide image intensifier viewed by a television system with a lead-oxide vidicon tube. The video signal is digitized to 8-bit accuracy every 100 ns and stored in any of three identical 256 by 256 random-access memories. Data may be temporally integrated to 13-bit gray-scale resolution. Several algorithms may be employed to process the data. All algorithms are presently dedicated hardware-based functions which were assembled in our laboratories. Although analog data may be processed from video tape, best results are achieved when image processing is performed on the original data. In the latter case the desired fully processed images appear during, or immediately after, the X-ray exposure. Present imaging modes include (1) three-spectrum K-edge imaging of iodine, (2) time-dependent subtraction angiography using intravenous injections of iodinated contrast agents, (3) time-dependent subtraction mask mode imaging of nonopacified hearts, (4) time-derivative displays of opacified and nonopacified hearts, and (5) cardiac phase displays of nonopacified hearts. K-edge subtraction imaging of opacified hearts at rates of 30 to 60 per second is also being investigated.