Abstract:
A method for measuring the apparent transverse relaxation time (“T* 2 ”) and apparent diffusion coefficient (“ADC”) of a hyperpolarized gas in a single breath-hold and consequently, with a single dose of the hyperpolarized gas contrast agent, is provided. The method employs a multiple-echo projection acquisition based pulse sequence. Individual images are reconstructed from data acquired during each of the individual echo times. Subsequently, T* 2  and ADC are calculated using these reconstructed images. Furthermore, the method produces images indicative of ADC that have isotropic resolution, allowing for more reliable image registration. The inter-echo spacing and diffusion weighting b-value are varied during the pulse sequence employed when practicing the present invention; thus, a significant separation between the effects of diffusion and T* 2  decay on the detected MR signals is possible. This separation allows for reliable measurements of these two parameters from a single echo-train.

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
     This application claims the benefit of U.S. Provisional patent application Ser. No. 61/046,817 filed on Apr. 22, 2008 and entitled “Method For Simultaneously Measuring T2* and Diffusion With Magnetic Resonance Imaging”. 
    
    
     STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH 
     This invention was made with United States Government support awarded by the following agency: NIH EB0002075. The United States Government has certain rights in this invention. 
    
    
     BACKGROUND OF THE INVENTION 
     The field of the invention is magnetic resonance imaging and systems. More particularly, the invention relates to methods for simultaneously measuring T* 2  and diffusion of a hyperpolarized gas contrast agent. 
     When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B 0 ), the individual magnetic moments of the nuclei in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B 1 ) that is in the x-y plane and that is near the Larmor frequency, the net aligned moment, M z , may be rotated, or “tipped”, into the x-y plane to produce a net transverse magnetic moment M xy . A signal is emitted by the excited nuclei or “spins”, after the excitation signal B 1  is terminated, and this signal may be received and processed to form an image. 
     When utilizing these “MR” signals to produce images, magnetic field gradients (G x , G y , and G z ) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received MR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques. 
     The measurement cycle used to acquire each MR signal is performed under the direction of a pulse sequence produced by a pulse sequencer. Clinically available MRI systems store a library of such pulse sequences that can be prescribed to meet the needs of many different clinical applications. Research MRI systems include a library of clinically proven pulse sequences and they also enable the development of new pulse sequences. 
     The MR signals acquired with an MRI system are signal samples of the subject of the examination in Fourier space, or what is often referred to in the art as “k-space”. Each MR measurement cycle, or pulse sequence, typically samples a portion of k-space along a sampling trajectory characteristic of that pulse sequence. Most pulse sequences sample k-space in a raster scan-like pattern sometimes referred to as a “spin-warp”, a “Fourier”, a “rectilinear”, or a “Cartesian” scan. The spin-warp scan technique employs a variable amplitude phase encoding magnetic field gradient pulse prior to the acquisition of MR spin-echo signals to phase encode spatial information in the direction of this gradient. In a two-dimensional implementation (“2DFT”), for example, spatial information is encoded in one direction by applying a phase encoding gradient, G y , along that direction, and then a spin-echo signal is acquired in the presence of a readout magnetic field gradient, G x , in a direction orthogonal to the phase encoding direction. The readout gradient present during the spin-echo acquisition encodes spatial information in the orthogonal direction. In a typical 2DFT pulse sequence, the magnitude of the phase encoding gradient pulse, G y , is incremented, ΔG y , in the sequence of measurement cycles, or “views” that are acquired during the scan to produce a set of k-space MR data from which an entire image can be reconstructed. 
     There are many other k-space sampling patterns used by MRI systems. These include “radial”, or “projection reconstruction” scans in which k-space is sampled as a set of radial sampling trajectories extending from the center of k-space. The pulse sequences for a radial scan are characterized by the lack of a phase encoding gradient and the presence of a readout gradient that changes direction from one pulse sequence view to the next. There are also many k-space sampling methods that are closely related to the radial scan and that sample along a curved k-space sampling trajectory rather than the straight line radial trajectory. 
     An image is reconstructed from the acquired k-space data by transforming the k-space data set to an image space data set. There are many different methods for performing this task and the method used is often determined by the technique used to acquire the k-space data. With a Cartesian grid of k-space data that results from a 2D or 3D spin-warp acquisition, for example, the most common reconstruction method used is an inverse Fourier transformation (“2DFT” or “3DFT”) along each of the 2 or 3 axes of the data set. With a radial k-space data set and its variations, the most common reconstruction method includes “regridding” the k-space samples to create a Cartesian grid of k-space samples and then perform a 2DFT or 3DFT on the regridded k-space data set. In the alternative, a set of radial k-space data can also be transformed to Radon space by performing a 1DFT of each radial projection view. 
     Certain noble gases can be put into a hyperpolarized state and employed as contrast agents in MR imaging applications, yielding substantial SNR increases over traditional proton MR imaging methods. Imaging methods that employ noble gases in the aforementioned manner are disclosed, for example, in U.S. Pat. No. 6,426,058. Of particular interest, is the use of hyperpolarized gas for imaging the air-filled spaces within the lung. In such an imaging study, a hyperpolarized noble gas such as helium ( 3 He) or xenon ( 129 Xe) is inhaled into the lungs prior to the MRI scan. While the spatial resolution attainable in MR images acquired from hyperpolarized gas studies is less than conventional MR imaging techniques, the sensitivity of MRI to the diffusion of hyperpolarized gases within the lung microstructure provides a mechanism for assessing the viability of lung tissue. Using diffusion weighted MRI (“DWI”), the apparent diffusion coefficient (“ADC”) of a hyperpolarized gas, such as helium-3, in the lung can be determined. 
     In DWI methods, motion sensitizing magnetic field gradients are applied so that the MR images include contrast related to the diffusion of water or other fluid molecules, such as hyperpolarized gas. By applying the diffusion gradients in selected directions during the MRI measurement cycle, diffusion weighted images are acquired from which the ADC is obtained for each voxel location in the reconstructed image. Hyperpolarized gas molecules diffuse less readily when they are restricted by the microstructure of the surrounding tissues. Hence, in diseases such as emphysema, which is characterized by a breakdown in the alveolar walls of the lung, measurements of the ADC of the inhaled hyperpolarized gas can be employed to assess tissue viability. Diffusion weighted MR imaging methods using hyperpolarized gas have been developed; however, the current techniques employ bipolar diffusion sensitizing gradients. Images acquired with and without these bipolar gradients present are used to determine the ADC of the gas in the lung tissues. 
     The apparent transverse relaxation, or T 2   * , for a proton species has also found use for assessing tissue viability. Mapping of T* 2  for a proton spin species has been demonstrated using multi-echo projection acquisition (“PR”) techniques. Also, a technique using multiple image acquisitions (each acquired in a new breath-hold and at a different echo time) has been used for T* 2  mapping in the lungs using PR methods. However, this method requires multiple breath-holds and the T* 2  in the lungs has been shown to be highly dependent on lung inflation volume and, therefore, repeatability between breaths. 
     It would therefore be desirable to provide a method that can simultaneously measure the diffusion and spin relaxation parameters of a hyperpolarized gas contrast agent. More particularly, it would be desirable to provide such a method that is applicable for a single dose of a hyperpolarized gas contrast agent and can be employed within a single breath-hold by the subject. 
     SUMMARY OF THE INVENTION 
     The present invention overcomes the drawbacks of previous methods by providing a method for imaging the apparent transverse relaxation time (“T* 2 ”) and apparent diffusion coefficient (“ADC”) of a hyperpolarized gas in a single breath-hold and consequently, with a single dose of the hyperpolarized gas contrast agent. 
     The present invention provides a method for simultaneously imaging T* 2  and the ADC in the lungs, or other airspace in the body, using a hyperpolarized gas contrast agent. The method employs a multiple-echo projection acquisition based pulse sequence. Individual images are reconstructed from data acquired during each of the individual echo times. The T* 2  and ADC are then calculated using these reconstructed images. Furthermore, the present invention provides a method for producing images indicative of ADC that have isotropic resolution, allowing for more reliable image registration. 
     It is an aspect of the invention to provide a method that accurately measures ADC and T* 2  simultaneously in one breath-hold. Because the inter-echo spacing, TE n , and diffusion weighting b-value, b n , are varied during the pulse sequence employed when practicing the present invention, a significant separation between the effects of diffusion and T* 2  decay on the detected MR signals is possible. This separation allows for reliable measurements of these two parameters from a single echo-train. In this manner, a method for accurately measuring ADC and T* 2  simultaneously in one breath-hold is provided. 
     The foregoing and other aspects and advantages of the invention will appear from the following description. In the description, reference is made to the accompanying drawings which form a part hereof, and in which there is shown by way of illustration a preferred embodiment of the invention. Such embodiment does not necessarily represent the full scope of the invention, however, and reference is made therefore to the claims and herein for interpreting the scope of the invention. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  is a block diagram of an MRI system that employs the present invention; 
         FIG. 2  is a graphic representation of an exemplary pulse sequence employed by the MRI system of  FIG. 1  when practicing an embodiment of the present invention; 
         FIG. 3  is a graphic representation of the radial trajectories of k-space samples acquired with the pulse sequence of  FIG. 2 ; and 
         FIG. 4  is a flowchart setting forth the steps of a method for simultaneously measuring the T* 2  and apparent diffusion coefficient (“ADC”) of a hyperpolarized gas. 
     
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     Referring particularly to  FIG. 1 , the preferred embodiment of the invention is employed in an MRI system. The MRI system includes a workstation  110  having a display  112  and a keyboard  114 . The workstation  110  includes a processor  116  that is a commercially available programmable machine running a commercially available operating system. The workstation  110  provides the operator interface that enables scan prescriptions to be entered into the MRI system. The workstation  110  is coupled to four servers: a pulse sequence server  118 ; a data acquisition server  120 ; a data processing server  122 , and a data store server  123 . The workstation  110  and each server  118 ,  120 ,  122  and  123  are connected to communicate with each other. 
     The pulse sequence server  118  functions in response to instructions downloaded from the workstation  110  to operate a gradient system  124  and an RF system  126 . Gradient waveforms necessary to perform the prescribed scan are produced and applied to the gradient system  124  that excites gradient coils in an assembly  128  to produce the magnetic field gradients G x , G y , and G x  used for position encoding MR signals. The gradient coil assembly  128  forms part of a magnet assembly  130  that includes a polarizing magnet  132  and a whole-body RF coil  134 . 
     RF excitation waveforms are applied to the RF coil  134  by the RF system  126  to perform the prescribed magnetic resonance pulse sequence. Responsive MR signals detected by the RF coil  134  or a separate local coil (not shown in  FIG. 1 ) are received by the RF system  126 , amplified, demodulated, filtered and digitized under direction of commands produced by the pulse sequence server  118 . The RF system  126  includes an RF transmitter for producing a wide variety of RF pulses used in MR pulse sequences. The RF transmitter is responsive to the scan prescription and direction from the pulse sequence server  118  to produce RF pulses of the desired frequency, phase and pulse amplitude waveform. The generated RF pulses may be applied to the whole body RF coil  134  or to one or more local coils or coil arrays (not shown in  FIG. 1 ). 
     The RF system  126  also includes one or more RF receiver channels. Each RF receiver channel includes an RF amplifier that amplifies the MR signal received by the coil to which it is connected and a detector that detects and digitizes the I and Q quadrature components of the received MR signal. The magnitude of the received MR signal may thus be determined at any sampled point by the square root of the sum of the squares of the I and Q components:
 
 M =√{square root over ( I   2   +Q   2 )},
 
     and the phase of the received MR signal may also be determined: 
     
       
         
           
             ϕ 
             = 
             
               
                 
                   tan 
                   
                     - 
                     1 
                   
                 
                 ⁡ 
                 
                   ( 
                   
                     Q 
                     I 
                   
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     The pulse sequence server  118  also optionally receives patient data from a physiological acquisition controller  136 . The controller  136  receives signals from a number of different sensors connected to the patient, such as ECG signals from electrodes or respiratory signals from a bellows. Such signals are typically used by the pulse sequence server  118  to synchronize, or “gate”, the performance of the scan with the subject&#39;s respiration or heart beat. 
     The pulse sequence server  118  also connects to a scan room interface circuit  138  that receives signals from various sensors associated with the condition of the patient and the magnet system. It is also through the scan room interface circuit  138  that a patient positioning system  140  receives commands to move the patient to desired positions during the scan. 
     The digitized MR signal samples produced by the RF system  126  are received by the data acquisition server  120 . The data acquisition server  120  operates in response to instructions downloaded from the workstation  110  to receive the real-time MR data and provide buffer storage such that no data is lost by data overrun. In some scans the data acquisition server  120  does little more than pass the acquired MR data to the data processor server  122 . However, in scans that require information derived from acquired MR data to control the further performance of the scan, the data acquisition server  120  is programmed to produce such information and convey it to the pulse sequence server  118 . For example, during prescans MR data is acquired and used to calibrate the pulse sequence performed by the pulse sequence server  118 . Also, navigator signals may be acquired during a scan and used to adjust RF or gradient system operating parameters or to control the view order in which k-space is sampled. And, the data acquisition server  120  may be employed to process MR signals used to detect the arrival of contrast agent in a magnetic resonance angiography (MRA) scan. In all these examples the data acquisition server  120  acquires MR data and processes it in real-time to produce information that is used to control the scan. 
     The data processing server  122  receives MR data from the data acquisition server  120  and processes it in accordance with instructions downloaded from the workstation  110 . Such processing may include, for example: Fourier transformation of raw k-space MR data to produce two or three-dimensional images; the application of filters to a reconstructed image; the performance of a backprojection image reconstruction of acquired MR data; the calculation of functional MR images; the calculation of motion or flow images, etc. 
     Images reconstructed by the data processing server  122  are conveyed back to the workstation  110  where they are stored. Real-time images are stored in a data base memory cache (not shown) from which they may be output to operator display  112  or a display  142  that is located near the magnet assembly  130  for use by attending physicians. Batch mode images or selected real time images are stored in a host database on disc storage  144 . When such images have been reconstructed and transferred to storage, the data processing server  122  notifies the data store server  123  on the workstation  110 . The workstation  110  may be used by an operator to archive the images, produce films, or send the images via a network to other facilities. 
     Referring particularly to  FIG. 2 , a three-dimensional (“3D”) projection reconstruction pulse sequence includes three readout gradient waveforms  206 ,  208 , and  210 . Each readout gradient waveform includes an initial dephasing lobe  212  followed by three readout gradient lobes  216 ,  220 , and  224 , and then a rephasing lobe  228 . First, a radiofrequency (“RF”) pulse  200  is played out in the presence of a slice-selective gradient  202 , such that transverse magnetization is produced in a prescribed imaging slice. The slice-selective gradient  202  includes a rephasing lobe  204  that acts to mitigate the effects of unwanted phase accruals during the application of the slice-selective gradient  202 . After the application of the RF excitation pulse  200 , k-space data is acquired during the entire playout of the readout gradient waveforms  206 ,  208 , and  210 . The RF excitation pulse  200  is, for example, a spectrally selective RF excitation pulse that is set to the resonance frequency of helium-3. However, it will be appreciated by those skilled in the art that the RF excitation pulse can be set to the resonance frequency of other spin species, such as that for xenon-129. A modulated flip angle is implemented to uniformly utilize the finite magnetization over each excitation such as, for example, the method described by G. W. Miller, et al., in “Hyperpolarized  3 He Lung Ventilation Imaging with B1-Inhomogeneity Correction in a Single Breath-Hold Scan,”  MAGMA,  2004; (16):218-226. 
     In the exemplary pulse sequence shown in  FIG. 2 , eight radial sampling trajectories are acquired during each repetition of the pulse sequence. Referring to  FIGS. 2 and 3 , during the application of dephasing lobes  212  k-space data is acquired by sampling a first half-echo  231  along a first radial trajectory  331 . This sampling is radially outward from the center of k-space and it is performed during the ramps and plateau of the dephasing lobe  212 . A small rotation gradient blip  214  is then applied to move the k-space sampling to another radial trajectory starting point as indicated by arrow  390 . During the first half of the readout gradient lobe  216  k-space data is acquired by sampling a second half-echo  232  along a second radial trajectory  332  back to the center of k-space. At the center of the readout gradient lobe  216  indicated by dotted line  217 , the level of each readout gradient G x , G y , and G z  is changed slightly to redirect sampling to another radial direction. As a result, during the playout of the remaining half of the readout gradient lobe  216 , k-space data is acquired by sampling a third half-echo  233  along a third radial sampling trajectory  333  which is directed away from the center of k-space. In this manner, k-space data is acquired during the playout of the entire readout lobe  216 , including its ramps. 
     Referring still to  FIGS. 2 and 3 , prior to playing out the second readout gradient lobes  220  another small rotation gradient blip  218  is applied to move the k-space sampling to another radial trajectory. As a result, when magnetic resonance (“MR”) signal acquisition is performed during the first half of readout lobe  220 , a fourth k-space sampling trajectory  334  is traversed back to the center of k-space such that a fourth half-echo  234  is sampled. Similar to the playing out of the first readout lobe,  216 , the level of each readout gradient G x , G y , and G z  is changed slightly to redirect sampling to another radial direction. As a result, during the playout of the remaining half of the readout gradient lobe  220 , k-space data is acquired by sampling a fifth half-echo  235  along a fifth radial sampling trajectory  335 , which is directed away from the center of k-space. Prior to playing out the last readout gradient lobes  224  another small rotation gradient blip  222  is applied to move the k-space sampling to another radial trajectory. As a result, when MR signal acquisition is performed during the first portion of readout lobe  224 , a sixth k-space sampling trajectory  336  is traversed back to the center of k-space such that a sixth half-echo  236  is sampled. 
     Again, the level of each readout gradient G x , G y , and G z  is changed slightly to redirect sampling to another radial direction in the remainder of the readout lobe  224  to redirect sampling to another radial direction. As a result, during the playout of the remaining portion of the readout gradient lobe  224 , k-space data is acquired by sampling a seventh half-echo  237  along a seventh radial sampling trajectory  337 , which is directed away from the center of k-space. Prior to playing out the dephasing gradient lobe  228  another small rotation gradient blip  226  is applied to move the k-space sampling to another radial trajectory. As a result, when MR signal acquisition is performed during the dephasing lobe  228 , an eighth k-space sampling trajectory  338  is traversed back to the center of k-space such that an eighth half-echo  238  is sampled. In particular, the first and second portions of the third readout gradient lobes  224  are not equal. In this manner, the second portion of the readout lobes  224  acts to impart a greater diffusion weighting to those MR signals acquired by sampling the eight half-echo  238 . The increased duration of the gradients results in a larger k-space sampling radius, as shown in  FIG. 3 . For example, image data acquired during the non-diffusion-weighted portion of the pulse sequence are acquired with projections in k-space having a radius of 64, while the diffusion weighted image data are acquired with projections in k-space having a radius of 256. 
     Following the sampling of the last half-echo  238 , a crusher gradient  240  is applied along each gradient axis to remove any unwanted magnetization so that signals in subsequent repetition time (“TR”) periods are not contaminated by residual magnetization. At the completion of the pulse sequence, therefore, a total of eight different radial trajectories in k-space capable of creating isotropic 256×256×256 pixel images are sampled during a repetition time (“TR”) period ranging, for example, from 3 to several milliseconds (“ms”). 
     Sampling during gradient ramping is performed to reduce the echo time and improve the overall data acquisition efficiency. To combine signals from ramp samples and multiple echoes effectively and robustly, the gradients are characterized using, for example, the method proposed by J. H. Duyn, et al., in “Simple Correction Method for k-Space Trajectory Deviations in MRI,”  JMR,  1998; (132):150-153. This characterization data is used to grid the acquired data at proper k-space locations. 
     Referring particularly now to  FIG. 4 , a method for simultaneously measuring the ADC and T* 2  of a hyperpolarized gas begins by administering the hyperpolarized gas contrast agent to a subject, as indicated in step  400 . Image data is subsequently acquired using, for example, the aforementioned pulse sequence, as indicated at step  402 . This acquired image, or k-space, data is then reconstructed to produce a series of images, as indicated at step  404 . In general, one image is reconstructed from each set of k-space data acquired for a given half-echo, with the earlier echoes more strongly weighted towards T* 2  decay. Therefore, and by way of example, eight images are reconstructed, one for each half-echo sampled by the aforementioned pulse sequence. 
     As discussed above, image data acquired during the non-diffusion-weighted portion of the pulse sequence is acquired with projections in k-space having a smaller sampling radius than the diffusion weighted image data. Thus, the diffusion weighted k-space data is regridded, as indicated at step  406 . This regridding essentially resamples the k-space data having a larger sampling radius to projection views of the smaller k-space radius. For example, k-space data from the seventh and eighth echoes,  237  and  238 , acquired using the modified eight half-echo pulse sequence above are gridded to the k-space radius of the first half-echo  231  (64 points). From the regridded k-space data, additional images are reconstructed, as indicated at step  408 . These images reconstructed from the regridded data therefore exhibit greater diffusion weighting, while minimizing T* 2  decay. Using these regridded images, along with the images reconstructed in step  404 , the apparent diffusion coefficient (“ADC”) and T* 2  of the hyperpolarized gas in the subject are estimated, as indicated at step  410 . A bi-exponential NMR signal model is employed to determine T* 2  and the apparent diffusion coefficient (“ADC”), D. This signal model has the form: 
     
       
         
           
             
               
                 
                   
                     
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     where S n  is the signal acquired from the n th  half-echo, S 0  is a signal acquired in the absence of diffusion weighting, D is the apparent diffusion coefficient, b n  is the b-value indicative of the amount of diffusion weighting applied during the n th  half-echo, TE n  is the inter-echo spacing for the n th  half-echo, and T* 2  is the apparent transverse relaxation time. Because of the differences in the signal decay in the diffusion weighted and T* 2  weighted sets of images, both ADC and T* 2  measures are simultaneously estimated by fitting the logarithm of the signal decay in these images, on a voxel-by-voxel basis, to a linearized form of the signal model in Eqn. (1). This linearized signal model has the form: 
     
       
         
           
             
               
                 
                   
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     As mentioned above, the linearized signal model is used to fit the logarithm of the signal decay, ln(S n /S 0 ), on a voxel-by-voxel basis. From this fit, T* 2  and D are determined. It should be appreciated by those skilled in the art that after T* 2  and D have been estimated on a voxel-by-voxel basis that an image indicative of these parameters can subsequently be produced. 
     The present invention has been described in terms of one or more preferred embodiments, and it should be appreciated that many equivalents, alternatives, variations, and modifications, aside from those expressly stated, are possible and within the scope of the invention.