Abstract:
A system and method for applying an RF excitation pulse to the region of interest (ROI) and a plurality of selective gradients to the ROI to elicit MR data pertaining to at least a first MR parameter from the ROI. The system and method also apply at least one diffusion gradient to the ROI to modulate the first MR parameter with a second MR parameter, acquire MR data from the ROI, and reconstruct a parametric map of the ROI using the MR data, wherein the parametric map is weighted based on the first MR parameter and modulated by the second MR parameter.

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
       [0001]    The present application is based on, claims the benefit of, and incorporates herein by reference U.S. Provisional Application Ser. No. 61/438,468, filed Feb. 1, 2011, and entitled, “SYSTEM AND METHOD FOR DIFFUSION-MODULATED RELAXATION MAGNETIC RESONANCE IMAGING.” 
     
    
     STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH 
       [0002]    N/A. 
       BACKGROUND OF THE INVENTION 
       [0003]    The field of the invention is magnetic resonance imaging (“MRI”) systems and methods. More particularly, the invention relates to systems and methods for deriving new clinically-useful information from a plurality of contrast mechanisms. 
         [0004]    When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B 0 ), the individual magnetic moments of the excited 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 B1) 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 t . 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. 
         [0005]    In MRI systems, the excited spins induce an oscillating sine wave signal in a receiving coil. The frequency of this signal is near the Larmor frequency, and its initial amplitude, A 0 , is determined by the magnitude of the transverse magnetic moment M t . The amplitude, A, of the emitted NMR signal decays in an exponential fashion with time, t. The decay constant 1/T*2 depends on the homogeneity of the magnetic field and on T 2 , which is referred to as the “spin-spin relaxation” constant, or the “transverse relaxation” constant. The T 2  constant is inversely proportional to the exponential rate at which the aligned precession of the spins would dephase after removal of the excitation signal B 1  in a perfectly homogeneous field. The practical value of the T 2  constant is that tissues have different T 2  values and this can be exploited as a means of enhancing the contrast between such tissues. 
         [0006]    Another important factor that contributes to the amplitude A of the NMR signal is referred to as the spin-lattice relaxation process that is characterized by the time constant T 1 . It describes the recovery of the net magnetic moment M to its equilibrium value along the axis of magnetic polarization (z). The T 1  time constant is longer than T 2 , much longer in most substances of medical interest. As with the T 2  constant, the difference in T 1  between tissues can be exploited to provide image contrast. 
         [0007]    Thus, images weighted based on the T 1  or T 2  time constants can be referred to as relaxation weighted imaging; however, a variety of other contrast mechanisms have also been developed. For example, a so-called diffusion weighted imaging (“DWI”) pulse sequence uses motion sensitizing magnetic field gradients to obtain images having contrast related to the diffusion of water or other fluid molecules. Specifically, a DWI pulse sequence applies diffusion sensitizing magnetic field gradients in selected directions during the MRI measurement cycle to obtain MR images that have an image contrast related to the diffusion of water or other fluid molecules that occurred during the application of the diffusion gradients. Using these DWI images, an apparent diffusion coefficient (“ADC”) may be calculated for each voxel location in the reconstructed images. 
         [0008]    Though the particular information sought in given clinical application may dictate a desired contrast mechanism (for example, T 1  weighting, T 2  weighting, diffusion weighting, perfusion imaging, and the like), it is well known that biological tissue is often heterogeneous and, therefore, has heterogeneous MR parameters, including T 1  relaxation times, T 2  relaxation times, diffusion coefficients, and magnetization transfer (MT), to name but a few. Accordingly, in many clinical settings, it would often be desirable to perform multiple MR acquisitions, each focusing on different contrast mechanisms, to ensure that the clinician is provided with a broad spectrum of information that, ideally, provides a full and accurate picture of the subject. 
         [0009]    For instance, during acute stroke, the diffusion of ischemic lesion decreases significantly. While for chronic stroke and tumor tissue, diffusion often is elevated due to edema and change in vasculature. As such, it is often desirable to perform multiple MR studies to acquire diffusion-weighted images, as well as other measurements to probe heterogeneous tissue damage. However, in many clinical settings and, particularly, when time is of the essence to diagnose and intervene to minimize the potential impact of the underlying conditions, it may be impractical to perform a large number of extended imaging studies. 
         [0010]    Furthermore, MRI is known to be susceptible to partial volume effects, due to limited spatial resolution. Such limitations may be particularly severe when considering complex pathophysiological changes during disease states. For instance, ischemic tissue has an elevated T 2  relaxation constant within hours after hypoperfusion. However, the diffusion rate of ischemic tissue has a complex pattern, whereby it initially decreases but then slowly recovers in about one week. Again, when looking to consider complex pathophysiological changes resulting during changes in disease states, it is desirable, yet not always cost effective to conduct a series of imaging studies spanning a variety of MR parameters. 
         [0011]    Accordingly, it would be desirable to have a system and method that provides a clinician with the ability to acquire information about a variety of contrast mechanisms and MR parameters without requiring lengthy and/or repetitive imaging studies. 
       SUMMARY OF THE INVENTION 
       [0012]    The present invention overcomes the aforementioned drawbacks by providing a system and method for acquiring information about multiple MR parameters, such as T 1  relaxation times or T 2  relaxation times and diffusion information, during a combined imaging acquisition. Specifically, the present invention provides an MR pulse sequence that sensitizes T 2  MRI acquisitions to diffusion parameters to generate diffusion modulated T 2 -weighted MR images and, more particularly, to quantify T 2  using multiple T 2 -weighted MR images having similar diffusion weighting. For example, by superimposing diffusion gradients upon spin echo T 2  MRI acquisitions, filtered parametric T 2  maps and other, multi-MR parameter weighted images can be generated. 
         [0013]    In accordance with one aspect of the invention, a magnetic resonance imaging (MRI) system is disclosed that includes a magnet system configured to generate a polarizing magnetic field about at least a region of interest (ROI) of a subject arranged in the MRI system and a plurality of gradient coils configured to apply a gradient field with respect to the polarizing magnetic field. The system also includes a radio frequency (RF) system configured to apply RF excitation fields to the subject and a acquire MR image data therefrom and a computer programmed to control the plurality of gradient coils and the RF system. The computer is programmed to cause the system to apply an RF excitation pulse to the ROI and a plurality of selective gradients to the ROI to elicit MR data pertaining to at least a first MR parameter from the ROI. The computer is further programmed to cause the system to apply at least one diffusion gradient to the ROI to modulate the first MR parameter with a second MR parameter, acquire MR data from the ROI, and reconstruct a parametric map of the ROI using the MR data, wherein the parametric map is weighted based on the first MR parameter and modulated by the second MR parameter. 
         [0014]    In accordance with another aspect of the present invention, a method for acquiring images of a region of interest (ROI) of a subject using a magnetic resonance imaging system is disclosed that includes applying an RF excitation pulse to the ROI in the presence of a first slice selective gradient. The method also includes applying at least one diffusion sensitizing gradient, applying a refocusing RF pulse to the ROI in the presence of a second slice selective gradient, and apply a readout gradient to acquire a echo signal from the ROI. The method then includes reconstructing a parametric map from the acquired echo signal that is weighted based on the a MR relaxation parameter and modulated by a diffusion parameter. 
         [0015]    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 
         [0016]      FIG. 1  is a block diagram of an MRI system that employs the present invention; 
           [0017]      FIG. 2  is a diagram of spin-echo pulse sequence; 
           [0018]      FIG. 3A  is a diagram of a multi-MR parameter pulse sequence in accordance with the present invention; 
           [0019]      FIG. 3B  is a diagram of another multi-MR parameter pulse sequence in accordance with the present invention; and 
           [0020]      FIGS. 4A-4D  are graphs for comparing the clinical data acquired using the multi-MR parameter pulse sequences of the present invention with data acquired using traditional imaging techniques. 
       
    
    
     DESCRIPTION OF THE INVENTION 
       [0021]    Referring to  FIG. 1 , an exemplary MRI system  100  for use with the present invention is illustrated. The MRI system  100  includes a workstation  102  having a display  104  and a keyboard  106 . The workstation  102  includes a processor  108 , such as a commercially available programmable machine running a commercially available operating system. The workstation  102  provides the operator interface that enables scan prescriptions to be entered into the MRI system  100 . The workstation  102  is coupled to four servers: a pulse sequence server  110 ; a data acquisition server  112 ; a data processing server  114 , and a data store server  116 . The workstation  102  and each server  110 ,  112 ,  114  and  116  are connected to communicate with each other. 
         [0022]    The pulse sequence server  110  functions in response to instructions downloaded from the workstation  102  to operate a gradient system  118  and a radiofrequency (“RF”) system  120 . Gradient waveforms necessary to perform the prescribed scan are produced and applied to the gradient system  118 , which excites gradient coils in an assembly  122  to produce the magnetic field gradients G x , G y , and G z  used for position encoding MR signals. The gradient coil assembly  122  forms part of a magnet assembly  124  that includes a polarizing magnet  126  and a whole-body RF coil  128 . 
         [0023]    RF excitation waveforms are applied to the RF coil  128 , or a separate local coil (not shown in  FIG. 1 ), by the RF system  120  to perform the prescribed magnetic resonance pulse sequence. Responsive MR signals detected by the RF coil  128 , or a separate local coil (not shown in  FIG. 1 ), are received by the RF system  120 , amplified, demodulated, filtered, and digitized under direction of commands produced by the pulse sequence server  110 . The RF system  120  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  110  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  128  or to one or more local coils or coil arrays (not shown in  FIG. 1 ). 
         [0024]    The RF system  120  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  128  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: 
         [0000]        M =√{square root over ( I   2   +Q   2 )}  Eqn. 1;
 
         [0025]    and the phase of the received MR signal may also be determined: 
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         [0026]    The pulse sequence server  110  also optionally receives patient data from a physiological acquisition controller  130 . The controller  130  receives signals from a number of different sensors connected to the patient, such as electrocardiograph (“ECG”) signals from electrodes, or respiratory signals from a bellows or other respiratory monitoring device. Such signals are typically used by the pulse sequence server  110  to synchronize, or “gate,” the performance of the scan with the subject&#39;s heart beat or respiration. 
         [0027]    The pulse sequence server  110  also connects to a scan room interface circuit  132  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  132  that a patient positioning system  134  receives commands to move the patient to desired positions during the scan. 
         [0028]    The digitized MR signal samples produced by the RF system  120  are received by the data acquisition server  112 . The data acquisition server  112  operates in response to instructions downloaded from the workstation  102  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  112  does little more than pass the acquired MR data to the data processor server  114 . However, in scans that require information derived from acquired MR data to control the further performance of the scan, the data acquisition server  112  is programmed to produce such information and convey it to the pulse sequence server  110 . For example, during prescans, MR data is acquired and used to calibrate the pulse sequence performed by the pulse sequence server  110 . Also, navigator signals may be acquired during a scan and used to adjust the operating parameters of the RF system  120  or the gradient system  118 , or to control the view order in which k-space is sampled. The data acquisition server  112  may also 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  112  acquires MR data and processes it in real-time to produce information that is used to control the scan. 
         [0029]    The data processing server  114  receives MR data from the data acquisition server  112  and processes it in accordance with instructions downloaded from the workstation  102 . 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 generation of functional MR images; and the calculation of motion or flow images. 
         [0030]    Images reconstructed by the data processing server  114  are conveyed back to the workstation  102  where they are stored. Real-time images are stored in a data base memory cache (not shown in  FIG. 1 ), from which they may be output to operator display  112  or a display  136  that is located near the magnet assembly  124  for use by attending physicians. Batch mode images or selected real time images are stored in a host database on disc storage  138 . When such images have been reconstructed and transferred to storage, the data processing server  114  notifies the data store server  116  on the workstation  102 . The workstation  102  may be used by an operator to archive the images, produce films, or send the images via a network to other facilities. 
         [0031]    As mentioned above, relaxation weighted and diffusion weighted MRI are commonly used imaging techniques that have become a vital clinical resource. For example, MRI imaging using such contrast mechanisms are commonly used for studying pathologies such as acute stroke and tumor, and hold great promise for non-invasive evaluation of prognosis and novel therapeutics. For instance, when considering ischemic tissue, it is known that the T 2  relaxation constant of ischemic tissues becomes elevated within hours after hypoperfusion, making T 2 -weighted MRI a widely used imaging technique for quantifying stroke outcome. In fact, T 2  or fluid attenuated inversion recovery (FLAIR) T 2  MRI is often applied to image stroke lesions, which is vital for the non-invasive evaluation of treatment outcome and assessment of novel therapeutic agents. In addition, diffusion MRI is also very sensitive to acute stroke, and can detect ischemic lesions immediately after hypoperfusion. As a result, accurate measurement of tissue ADC and T 2  are important for properly characterizing stroke pathophysiology. 
         [0032]    Although extremely versatile and useful, MRI, however, may be subject to non-negligible partial volume effects due to its limited spatial resolution and slice thickness. Moreover, cerebral tissue is heterogeneous; it contains white matter (WM), gray matter (GM), cerebral spinal fluid (CSF) and blood vessels whose diffusion and relaxations properties may be drastically different upon ischemia. For example, normal cerebral tissue ADC is approximately 0.8 μm 2 /ms, while it decreases significantly to about 0.5 μm 2 /ms upon ischemia. In contrast, within hours of ischemia, ischemic tissue T 2  can elevate from its normal value of 60 ms to 100 ms and above, due to edema. Hence, if partial volume effect (in plane and slice thickness) is not properly accounted for, MRI measurement only reflects an ensemble average of normal and ischemic tissue properties, and may not be specific to stroke lesion. As a result, there may be severe mischaracterization of stroke prognosis, impeding automated image analysis and segmentation. 
         [0033]    To address these clinical needs, the present invention provides systems and methods for using the above-described systems to acquire multi-MR parameter images using a combined pulse sequence. For example, in accordance with the present invention, diffusion weighting may be applied to modulate T 2  or T 1  imaging. More particularly, diffusion weighting may be applied to modulate the standard spin echo T 2  MRI. 
         [0034]    Referring to  FIG. 2 , a pulse sequence diagram  200  for a spin-echo pulse sequence is illustrated. As illustrated, the spin-echo pulse sequence  200  includes an RF excitation pulse  202  that is played out in the presence of a slice selective gradient  204 . To mitigate signal losses resulting from phase dispersions produced by the slice selective gradient  204 , a rephasing lobe  206  is applied after the slice selective gradient  204 . Next, a refocusing RF pulse  212  is applied following a phase encoding gradient  208  and associated readout gradient  210 . In order to substantially reduce unwanted phase dispersions, along with the refocusing pulse  212 , a first crusher gradient  216  bridges the slice selective gradient  214  with a second crusher gradient  218 . A spin-echo MR signal  220  is acquired during the application of a readout gradient  222 . As is known in the art, the pulse sequence  200  may be repeated a plurality of times while stepping the phase encoding gradient  208  through a plurality of different values. This process may then be repeated with different slice selective gradients  204 ,  214  so as to acquire image data from different slice locations. Accordingly, the TR is defined as the time between RF excitation pulses  202  and the TE is the time between the RF excitation pulse  202  and the spin echo  220 . 
         [0035]    For the conventional T 2  MRI sequence described above, the MR signal at TE is given as: 
         [0000]        M ( TE )= M   SS   ·e   −TE/T     2   ·sin α  Eqn: 3;
 
         [0036]    where M SS  is the steady state magnetization, T 2  is the transverse relaxation time, and a is the excitation angle. The steady-state magnetization depends on several parameters including repetition time (TR), excitation angle a, longitudinal relaxation time (T 1 ), and spin density (M 0 ). For simplicity in discussion, it can be assumed that the TR is very long compared with the duration of the T 1  relaxation time, while the excitation angle is an ideal π/2 pulse. For cerebral tissue, it can be assumed that the fraction of normal and ischemic tissue is denoted by f n  and f i , respectively, and f n +f i =1. Also, for cerebral tissue, the T 2  relaxation time and diffusion rates are given by T 2n,i  and D n,i , respectively. With this background in place, MR signal from the conventional spin echo T 2  MRI sequence is given by: 
         [0000]    
       
         
           
             
               
                 
                   
                     
                       
                         
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         [0037]    As such, the ratio of ischemic tissue signal to total MR signal, that is, M i /(M i +M n )) can be shown to be: 
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         [0038]    Thus, it can be seen that the relative contribution from ischemic tissue will be enhanced at longer echo time. 
         [0039]    As can be seen, with conventional T 2 -weighted MRI, calibration of parametric T 2 , often with a simplistic assumption of mono-exponential decay, at best, can be achieved by varying echo time (TE). However, as will be described, the present invention provides a system and method for acquiring information about multiple MR parameters, such as T 1  relaxation times or T 2  relaxation times and diffusion information, during a combined imaging acquisition. Specifically, the present invention provides a system and method to generate multi-MR parameter weighted images. For example, in one particular application, the present invention provides an MR pulse sequence that sensitizes T 2  MRI acquisitions to diffusion parameters to generate parametric diffusion modulated T 2 -weighted MR images. 
         [0040]    Turning now to  FIG. 3A , an example of a multi-MR parameter pulse sequence  300  for acquiring data for multi-MR parameter weighted images is provided. The, exemplary, illustrated multi-MR parameter pulse sequence  300  includes diffusion sensitizing gradients  302  that are superimposed upon a spin echo (SE) T 2  sequence  304 . However, as will be described, it is contemplated that other multi-MR parameter pulse sequences, such as those that include diffusion sensitizing gradients  302  that are superimposed with other T 2  sequences or superimposed with T 1  sequences, may also be used in accordance with the present invention. 
         [0041]    In particular, the illustrated multi-MR parameter pulse sequence  300 , like a traditional SE pulse sequence, includes an RF excitation pulse  306  that is played out in the presence of a slice selective gradient  308 . To mitigate signal losses resulting from phase dispersions produced by the slice selective gradient  308 , a rephasing lobe  310  is applied after the slice selective gradient  308 . A refocusing RF pulse  312  is applied and, in order to substantially reduce unwanted phase dispersions, a first crusher gradient  314  bridges the slice selective gradient  316  with a second crusher gradient  318 . A spin-echo MR signal  320  is acquired during the application of a readout gradient. It is noted that an echo planner imaging (EPI) readout may be used for image readout, for example, so T 2  measures can be obtained with a single echo technique. As mentioned above, unlike traditional SE pulse sequences or traditional pulse sequences designed for diffusion-weighted imaging, the multi-MR parameter pulse sequence  300  includes diffusion sensitizing gradients  302  that are superimposed upon a SE T 2  sequence  304 . To this end, a pair of bipolar gradients  322 ,  324  are inserted before and after the refocusing RF pulse  312 . The inclusion of the pair of bipolar gradients  322 ,  324  defines two new quantities,  T ′ and δ T ′, that are subcomponents of  T . As will be described in further detail, using the present invention, the TE can be adjusted by varying δ T ′, while the magnitude and duration of diffusion gradients  302  are fixed because  T=T ′+δ T′.    
         [0042]    However, as illustrated in  FIG. 3B , it is contemplated that the diffusion modulated T 2  MRI sequence of  FIG. 3A , in which the variable delays are positioned before and after the pair of diffusion sensitizing gradients  322 ,  324 , may be modified. Specifically, as illustrated in  FIG. 3B , another configuration of a multi-MR parameter pulse sequence  350  is provided whereby variations of the TE do not affect the diffusion b-factor, thereby, allowing calibration of absolute T 2 . 
         [0043]    The, exemplary, illustrated multi-MR parameter pulse sequence  350  again includes diffusion sensitizing gradients  352  that are superimposed upon a spin echo (SE) T 2  sequence  354 . However, unlike the multi-MR parameter pulse sequence  300  of  FIG. 3A , in the multi-MR parameter pulse sequence  350  illustrated in  FIG. 3B , the variable delays, δ T ′, are positioned before and after the diffusion sensitizing gradients  352 . 
         [0044]    Again, like a traditional SE pulse sequence, the multi-MR parameter pulse sequence  350  includes an RF excitation pulse  306  that is played out in the presence of a slice selective gradient  358 . To mitigate signal losses resulting from phase dispersions produced by the slice selective gradient  358 , a rephasing lobe  360  is applied after the slice selective gradient  358 . A refocusing RF pulse  362  is applied and, in order to substantially reduce unwanted phase dispersions, a first crusher gradient  364  bridges the slice selective gradient  366  with a second crusher gradient  368 . A spin-echo MR signal  370  is acquired during the application of a readout gradient. 
         [0045]    As noted previously, however, unlike traditional SE pulse sequences or traditional pulse sequences designed for diffusion-weighted imaging, the multi-MR parameter pulse sequence  350  includes diffusion sensitizing gradients  352  that are superimposed upon a SE T 2  sequence  354 . To this end, diffusion gradients  372 ,  374  are inserted before and after the refocusing RF pulse  362 . In this configuration, the TE can be adjusted by varying δ T ′, while the magnitude and duration of diffusion gradients  372  are fixed because  T=T ′+δ T ′. However, variations of TE do not affect diffusion b-factor, allowing calibration of absolute T 2 . Based on conventional spin echo DWI, additional adjustment of delays can be put before and after diffusion gradients to vary the echo time. As such, a pre-determined diffusion b-value can be obtained for a serial echo time, allowing diffusion modulated T 2  MRI. 
         [0046]    Given that diffusion in, for example, cerebral tissue is anisotropic, isotropic diffusion weighting may be preferred to mitigate such effects. By simultaneously applying gradients along multiple directions, it is also efficient to reach a given moderate diffusion b-factor. This can be achieved by averaging multiple images with diffusion gradients applied along orthogonal directions, or several single-shot isotropic diffusion imaging techniques can be easily implemented. While on the other hand, such isotropic diffusion weighting is not necessary when parametric T 2  is calibrated. In addition, because the TE is prolonged due to diffusion gradients, the minimal echo time for the proposed multi-MR parameter pulse sequence  300 ,  350  may be slightly longer than that of the conventional SE T 2  sequence. However, for a representative diffusion b-factor of 500 mm 2 /s, the diffusion module duration is less than 15 ms with a gradient strength of 30 g/cm in a single shot trace diffusion MRI sequence. 
         [0047]    The above-described systems and methods can be applied to a variety of clinical settings to image a variety of MR parameters. For example, though the multi-MR parameter pulse sequences described above involve the superimposition of diffusion gradients over T 2 -weighted pulse sequences, the present invention may likewise be applied to, for example, superimposition of diffusion gradients over T 1  weighted pulse sequences. In particular, a clinician may chose to create a diffusion modulated T 1  relaxation image or a diffusion modulated T 1  relaxation image, or another multi-MR parameter weighted image, based on the clinical needs and indications. The present invention allows the clinician to create diffusion-modulated relaxation parametric MR images, for example, to quantify T 2  using multiple T 2 -weighted MR images having similar diffusion weighting and, thus, are substantially more sensitive than conventional T 2 -weighted MR images. For example, diffusion modulated T 2  parametric images would be advantageous in renal applications and neurological applications, such as stroke. Also, in the case of stroke, data suggests an early T 1  change that could likewise be reflected in diffusion modulated T 1  relaxation images. 
         [0048]    Exemplary Clinical Application: Stroke 
         [0049]    In the example of diffusion modulated T 2  imaging, the difference in diffusion and relaxation rates between ischemic and normal cerebral tissue may be more readily discerned. The rationale is that in comparison with ischemic lesion, normal cerebral tissue has higher diffusivity, hence, diffusion weighting preferentially attenuates normal tissue signal with respect to that from ischemic lesion. As such, MR measurements will be weighted toward ischemic tissue due to reduced diffusion. In addition, application of diffusion gradients will prolong the echo time, which will concomitantly suppress normal tissue MR signal. In fact, it has been noted that diffusion and T 2  relaxation contrasts may be closely coupled in stroke imaging. Particularly, it is known that hyperintensity in diffusion-weighted MRI (DWI) may be attributed not only to decreased diffusivity of ischemic tissue, but also to its elevated T 2 , and dubbed T 2  shine-through effect. In fact, calibration of quantitative ischemic tissue ADC is recommended to minimize the T 2  shine-through effects. Hence, there has been a long standing clinical need to expedite the acquisition of and increase the availability of both T 2  and diffusion information in the case of stroke, to name but one of many clinical applications. 
         [0050]    Hence, the present invention can utilize the reduction of diffusivity in ischemic tissue to relatively suppress MR signal contribution from normal tissue, and make the “ensemble MR” (diffusion modulated T 2 ) measurement more specific to ischemic lesion. In consistent with the notion of T 2  shine-through effect, the multi-MR parameter pulse sequence, such as the diffusion modulated T 2  MRI sequence, of the present invention can be used to enhance ischemic lesion based on its decreased diffusivity, which may be regarded as an inverse diffusion shine-through contrast. Similar as conventional parametric T 2  imaging, the proposed diffusion modulated T 2  MRI also requires pseudo-linear fitting of images obtained at multiple echo times in order to derive the parametric T 2 . It is important to point out that because identical diffusion weighting is applied for spin echo images, the obtained T 2  map is the average of T 2  value from two magnetization pools modulated by their diffusion property. As such, contribution from each component can be adjusted simply by varying the magnitude of diffusion b value. 
         [0051]    As illustrated above, when using the conventional spin echo T 2  MRI sequence to perform clinical studies of stroke, the relative contribution from ischemic tissue will be enhanced at longer echo time. On the other hand, when performing multi-MR parameter imaging in accordance with the present invention, such as using a diffusion-modulated T 2  MRI sequence, the MR signal acquired is given by: 
         [0000]        M ( TE )= m   0 ·( f   i   ·e   −b·D     i     ·e   −TE/T     2i   +(1 −f   i )· e   −b·D     n     ·e   −TE/T     2n   )   Eqn. 6;
 
         [0052]    where b is diffusion b value. The ratio of relative MR signal can be shown to be: 
         [0000]    
       
         
           
             
               
                 
                   
                     
                       
                         
                           
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         [0053]    In comparison with conventional T 2  MRI sequence, the contribution from normal tissue is attenuated due to diffusion modulation. It is important to note η′=η when b=0, as expected. 
         [0054]    To evaluate the performance of the diffusion modulated T 2  MRI technique of the present invention in the clinical application of stroke, the effect of diffusion and echo time on MR relaxation measurement may be evaluated using a simplistic 2-compartment tissue model (i.e., normal vs. ischemic tissue). In vivo results showed that ischemic tissue T 2  measured using the present invention is significantly higher than conventional spin echo T 2  measurement. In addition, the present invention can capture ischemic lesions not easily observable using the conventional T 2  MRI, strongly suggesting that the enhanced specificity to ischemic tissue may permit improved diagnosis. 
         [0055]    More particularly, numerical simulation can be applied to elucidate MR signal for both the conventional T 2  and the diffusion modulated T 2  MRI sequences of the present invention. Representative T 2  value of 60 and 100 ms, and diffusion coefficient of 0.8 and 0.5 μm 2 /ms were used for normal and ischemic tissue, respectively. Referring to  FIG. 4A , normalized MR signal from ischemic (solid) and normal tissue (dotted) with TE from 30 to 80 ms is illustrated. For illustration, assuming that 50 percent of cerebral tissue is ischemic, the contribution from ischemic tissue signal to the total MR signal (η) is a function of echo time. It increases with echo time, consistent with the fact that ischemic tissue T 2  is elevated from that of normal tissue. Fitting the MRI signal against echo time with a commonly used single exponential decay function, the experimentally measured T 2  depends on ratio of ischemic tissue (f), as shown in  FIG. 4B . The measured T 2  is an ensemble average of ischemic and normal tissue T 2 , weighted by their fraction concentration. In contrast, diffusion modulated T 2  MRI data acquired in accordance with the present invention has additional contrast based on tissue diffusivity. For simulation, the echo time is increased by 20 ms to take into account of diffusion gradients.  FIG. 4C  shows that the ischemic tissue MR signal contribution (η′) is significantly enhanced from that without diffusion weighting ( FIG. 4C  vs.  FIG. 4A ), even at a moderate b value of 500 s/mm 2 . In addition, the experimentally derived relaxation time varies not only with ischemic tissue fraction, but also depends on the applied diffusion b-value as it modulates the relative signal contribution from ischemic and normal tissues.  FIG. 4D  shows that T 2  obtained from diffusion modulated MR sequence is higher than that obtained with conventional MR sequence (dashed line). In addition, the difference between T 2  measurements increases at large b value. 
         [0056]    T 2  maps obtained with the diffusion modulated spin echo MRI pulse sequence of the present invention clearly detected ischemic infarction in consecutive slices. In addition, the obtained T 2  value is significantly higher than that obtained with conventional spin echo MRI, suggesting that measurement from the diffusion modulated T 2  MRI is more specific to ischemic tissue. Using this phenomenon, automated and threshold-based lesion detection algorithms may be readily applied using the present invention with accuracy previously unachieved. 
         [0057]    Therefore, the present invention illustrates that parametric T 2  maps obtained from diffusion modulated T 2  MRI sequences allow detection of ischemic lesions not easily observable using conventional spin echo T 2  MRI. In addition, MR identified ischemic lesion can be confirmed by histology. With this in mind, the present invention may be used to allow early and accurate assessment of ischemic lesion. In fact, it has been shown that MCAO ischemic tissue damage is still evolving beyond 24 hours, and it would be desirable to develop a T 2  mapping technique that is more specific to ischemic tissue, allowing sensitive and early calibration of infarction. While T 2  MRI has been commonly used to assess stroke lesion for permanent and complete MCAO, ischemic tissue outcome may be very heterogeneous for transient stroke. 
         [0058]    Reperfusion, DWI, T 2  Lesion, and Biological Significance 
         [0059]    As described above, the improved detection of reperfusion using the present invention is attributed to the fact that diffusion weighting causes additional modulation of normal tissue MR signal, hence, making final MR measurements more specific to ischemic tissue. It is also important to point out that the diffusion-modulated T 2  MRI suppresses not only normal cerebral tissue, but also other cerebral components of high diffusivity/pseudo-diffusivity, such as cerebral spinal fluid (CSF) and vascular blood signal. To this end, it may be, in certain clinical settings, desirable to use single shot isotropic diffusion sensitizing gradients to minimize artifacts due to anisotropic diffusion in cerebral tissue. However, it is worthwhile to note many other gradient forms can also be utilized, such as that proposed by Song et al. In addition, given that a parametric T 2  map is obtained, isotropic diffusion weighting is not required in many situations, as long as adequate diffusion weighting is obtained. 
         [0060]    While the pulse sequences of the present invention provide enhanced detectability of ischemic lesion, it can be argued that similar detectability may be obtained using the conventional T 2  MRI, provided much finer spatial resolution. Although this may be the case, it remains advantageous to utilize the pulse sequences of the present invention when scan time, signal to noise ratio (SNR) and hardware limitations are considered. In addition, although the diffusion modulated T 2  MRI of the present invention is a hybrid sequence of diffusion weighted MRI and spin echo T 2  MRI, its contrast is unlikely simple replication of conventional T 2  and diffusion lesion. It is so because it combines diffusion weighting and spin echo T 2  MRI of serial echo time, and the obtained parametric T 2  map correlates diffusion and relaxation properties of ischemic tissue. As such, diffusion modulated T 2  MRI provides complementary information to both spin echo T 2  MRI and diffusion MRI. 
         [0061]    Further still, it is contemplated that diffusion-modulated T 2  MRI in accordance with the present invention may be extended to T 2 * measurement. Diffusion weighting may attenuate MR signal from normal tissue during acute stroke and, hence, makes both parametric T 2  and T 2 * more specific toward ischemic tissue. As such, it may allow more accurate estimation of tissue oxygen extraction ratio (OER), which is particularly important for elucidating the oxygen metabolic status of perfusion/diffusion lesion mismatch. In fact, surrogate imaging markers of ischemic physilogical status such as tissue pH and oxygen metabolism may provide vial and complementary information to commonly used perfusion and diffusion MRI, and ultimately, may facilitate development and evaluation of novel therapeutic agents and treatments. 
         [0062]    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.