Magnetic resonance imaging (MRI) is one of the most versatile and fastest growing modalities in medical imaging. Since the discovery by Dr. Raymond Damadian in the early 1970s that nuclear magnetic resonance techniques can be used to scan the human body to yield useful diagnostic information, medical NMR imaging devices have been developed for obtaining NMR images of the internal structures of patients. Subsequently, much effort has been expended to improve and refine the techniques used for obtaining NMR images, as well as to determine the diagnostic utility of NMR images. As a result, NMR imaging, or magnetic resonance imaging, as it is sometimes known, has today proven to be an extremely useful tool in the medical community for the purposes of detecting and diagnosing abnormalities within the body.
Conventional magnetic resonance imaging techniques generally utilize pulsed magnetic field gradients to spatially encode NMR signals from various portions of an anatomical region of interest. The pulsed magnetic field gradients, together with radio frequency excitation of the nuclear spins and acquisition of signal information, are commonly referred to as a pulse sequence.
The basic science behind NMR imaging is as follows. Pulsed current through a set of conductors will produce a magnetic field external to the conductors; the magnetic field generally has the same time course of development as the current in the conductors. The conductors may be distributed in space to produce three orthogonal gradients: X, Y, and Z. Each of the gradients may be independently pulsed by a separate time-dependent current waveform.
In order to construct images from the collection of NMR signals, conventional NMR imaging equipment generally utilizes magnetic field gradients for selecting a particular slice (volume) or plane of the object to be imaged and for encoding spatial information into the NMR signals. For example, one conventional technique involves subjecting an object to a continuous static homogenous field extending along a first direction, and to sets of sequences of orthogonal magnetic field gradients. Each set of orthogonal magnetic field gradient sequences generates a magnetic field component in the same direction as the static field, but the gradients have strengths that vary along their directions.
Generally, the NMR phenomenon occurs in atomic nuclei having an odd number of protons and/or neutrons. Due to the spins of the protons and neutrons, each such nucleus exhibits a magnetic moment. As a result, when a sample composed of such nuclei is placed in the homogeneous magnetic field, a greater number of nuclear magnetic moments align with the direction of the magnetic field to produce a net macroscopic magnetization in the direction of the field. Under the influence of the magnetic field, the magnetic moments precess about the axis of the field at a frequency that is dependent upon the strength of the applied magnetic field and on the characteristics of the nuclei. The angular precession frequency, ω, also referred to as the Larmor frequency, is given by the equation ω=γB, where γ is the gyro-magnetic ratio (which is a constant for each particular atomic nucleus) and B is the strength of the magnetic field acting upon the nuclear spins.
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but the randomly-oriented magnetic components in the perpendicular (transverse) plane (X-Y plane) cancel one another. If, however, the tissue is also subjected to a magnetic field (excitation field B1) that is in the X-Y plane and that is at the Larmor frequency, the net aligned moment, M, may be rotated, or “tipped”, into the X-Y plane to produce a net transverse magnetic moment Mt. Mt rotates, or spins, in the X-Y plane at the Larmor frequency. The practical value of this phenomenon resides in the signal that is emitted by the excited spins after the excitation field B1 is removed.
Thus, the orientation of the moment (also known as the magnetization) M can be perturbed by the application of a magnetic field oscillating at the Larmor frequency, which has the effect of rotating the magnetization away from the direction of the static field. Typically, the oscillating magnetic field is applied in a direction orthogonal to the direction of the static magnetic field by means of a radio frequency (RF) pulse through coils connected to a radio frequency transmitting apparatus. In essence, the net magnetic vector or orientation of magnetization M is rotated away from the direction of the static field. One typical RF pulse is that which has either sufficient magnitude or duration to rotate the magnetization M into a transverse plane (that is, 90 degrees from the direction of the static field) and thus is known as a 90-degree RF pulse. Similarly, if the magnitude or duration of the RF pulse is selected to be twice that of a 90-degree pulse, the magnetization M will change direction 180 degrees from the main or static magnetic field, and the excitation pulse is called a 180-degree RF pulse.
Accordingly, a typical imaging procedure involves the use of three orthogonal magnetic field gradients, X, Y, and Z, which are pulsed coordinately along with bursts of radio frequency energy. For example, the Z gradient is pulsed on for two brief time periods. A 90-degree radio frequency pulse in the first time period and a 180-degree radio frequency pulse in the second time period are used to select a slice of the anatomy of interest, and to induce the nuclear spin system within that slice to generate an NMR signal. Once the slice is selected by the Z gradient, the two remaining orthogonal gradients are used to confer spatial encoding on the NMR signal in the two orthogonal directions. For example, the Y gradient will encode on the basis of phase advances imparted on a series of signal responses by using a pulsed gradient waveform of progressively increasing area. The X gradient, which is pulsed on during the signal collection period, will frequency-encode the NMR signal in the third orthogonal direction.
When the RF excitation pulse is stopped (by turning the RF transmitter off), the nuclear spins tend to slowly realign or relax back to the equilibrium position. At this time, the spins emit an NMR signal, which can be detected with an RF receiver coil (which may be, and often is, the same coil as that used with the transmitter). The emitted NMR signal is dependent on three basic parameters, namely, the density of the excited nuclei, the spin-lattice (longitudinal) relaxation time (T1), and the spin-spin (transverse) relaxation time (T2). The latter two parameters are both exponential time constants that characterize the rate of return to equilibrium of the longitudinal and transverse magnetization components following the application of the perturbing RF pulse. These NMR parameters of spin density, T1 and T2, are related to the atomic nuclei subjected to the NMR phenomenon.
In accordance with this technique, nuclear spins in a selected plane are excited by a selective RF pulse, in the presence of one of the magnetic field gradients. The frequency of the selective RF pulse corresponds to the Larmor frequency for only the selected plane of the object as determined by the magnetic field gradient imposed on the static magnetic field. The applied magnetic field gradient is designated as the slice-selection gradient. The selected plane will therefore extend in a direction perpendicular to the gradient direction of the slice-selection magnetic field gradient. The selected spins are then subjected to the other magnetic field gradients (which can be designated as the read-out and phase-encoding magnetic field gradients). A plurality of repetitions are utilized in which the amplitude of the phase-encoding gradient is varied for each repetition and in which the read-out gradient is applied during the reading out of the generated NMR signals.
The NMR signal is processed to yield images that give an accurate representation of the anatomical features in the selected slice, as well as provide excellent soft tissue contrast. NMR signals may be processed using various algorithms, depending upon the precise nature of the data acquisition procedure. However, all methods employed rely on the ability to spatially encode the signal information by making use of the magnetic field gradients, which are time modulated and sequentially pulsed in various modes to effect the desired result.
For example, the received NMR signals may be transformed by utilizing, for example, conventional two-dimensional Fourier transform techniques. The read out magnetic field and phase-encoding magnetic field gradients encode spatial information into the collection of NMR signals so that two-dimensional images of the NMR signals in the selected plane can be constructed. During the scanning sequence, the various magnetic field gradients are repeatedly switched on and off at the desired intervals.
Many NMR imaging schemes rely on a collection of spin-echo NMR signals. In utilizing spin-echo signals, a 90-degree RF excitation pulse is followed by the application of a 180-degree rephasing RF pulse at a predetermined time interval after the 90-degree pulse. This produces a spin-echo signal at a corresponding time interval after the application of the 180-degree RF pulse. In NMR parlance, the time that the spin-echo NMR signal is produced after the 90-degree RF excitation pulse is designated as TE (for time of echo). Thus, the 180-degree RF pulse is applied at a time interval of TE/2 after the 90-degree RF pulse.
In the application of NMR principles to medicine and medical diagnostic imaging of live human subjects, NMR signals are obtained for a multitude of small areas in a patient, known as picture elements or pixels. The pixels are used to construct an image or pictorial representation of a particular area of the patient being examined. More particularly, the intensity of the NMR signals is measured for the multitude of pixels. The intensity of each signal is a complex function of the tissue-related parameters used in gathering the image information.
For example, it is known that variations in the relaxation times T1 and T2 are closely associated with the differences between healthy and diseased tissue, and thus, from a diagnostic viewpoint, images that display or show significant T1 and/or T2 contrast have proven to be of great diagnostic interest. Unfortunately, in conventional techniques for obtaining both T1-weighted data and T2-weighted data, not only are separate scans required for obtaining T1-weighted and T2-weighted images, but further, additional T1-weighted scans are required to obtain T1-weighted images for the same number of slices for which T2-weighted scans are obtained. This increases the number of scans of the patient that must be performed and the time necessary to complete such scans. After each imaging scan that is performed on a patient, it is generally necessary to allow the patient to rest. Also, a certain amount of time is necessary when conducting a scan for operator setup, loading information into the apparatus with respect to the conditions and sequencing for collection of data, etc. Therefore, in order to obtain, using conventional techniques, T2-weighted images for a plurality of planes and a corresponding number of T1-weighted images, the total scanning time is quite long. It is apparent that one of the major problems with medical NMR imaging is patient throughput. Thus, numerous efforts have been devoted to the development of techniques for obtaining images in a shorter period of time.
Although various efforts have been devoted to the development of techniques for shortening the scan times, to date, they have generally resulted in a sacrifice of the diagnostic quality of the information obtained, and thus, have not yet proven satisfactory.
One factor contributing to the length of an imaging time period is the period of time required for the return of the nuclear magnetizations to equilibrium prior to the subsequent excitation. A method that has been used to shorten this time period is known as the driven equilibrium pulse technique. In typical driven equilibrium techniques utilized with spin-echo sequences, spins in the X-Y plane are driven back to alignment with the Z-axis in order to shorten the time period required for the spins to return to equilibrium. As a result, the data acquisition time is shortened, and image contrasts may be manipulated in pulse sequences at a high repetition rate.
However, driven equilibrium techniques have not become a standard on clinical scanners. A major problem with conventional driven equilibrium techniques is the resulting lack of contrast control in the derived image. Some of the difficulty may be related to eddy current control and RF phase control in MRI scanners. That is, in order to obtain accurate diagnostic information using the conventional driven equilibrium technique, it is important that all the spins in the selected slice are precisely in phase, and that the 90-degree Z-restoring pulse is delivered along an axis that is exactly perpendicular to the direction along which the transverse magnetization is focused. Even small deviations in these parameters caused by eddy currents generated by the gradients are enough to seriously degrade the amount of magnetization restored to the Z-axis.
Another technique for decreasing the scan and image capture time of an NMR imaging procedure is the use of fast-spin echo methods. Such methods involve the acquisition of multiple spin-echo signals from a single excitation pulse in which each acquired echo signal is separately phase-encoded. Each pulse sequence therefore results in the acquisition of a plurality of views, and a plurality of pulse sequences is typically utilized to acquire a complete set of image data. However, using such techniques, although the TE time interval can be varied, the repetition rate is the same for each produced image.
In an effort to improve scanning throughput, most conventional scan methodologies utilize multiple-slice techniques. In a multiple-slice scan procedure, a patient must remain stationary throughout a prolonged scan process, during which an entire sequence of slices is captured. Patients often get uncomfortable during this process, and may fidget or move involuntarily. These movements, however small, create motion artifacts that make diagnosis difficult, or even impossible. This effect is particularly significant in a multi-slice process, because if a patient moves at any time, all the slices will show the motion artifacts and the entire scan procedure will have to be repeated.
This effect is more prevalent when performing a standing MRI scan, during which the patient is asked to stand during the scanning procedure. Images show dramatic differences in anatomy, due to the different weight-bearing qualities present in the human body, when a person is standing compared to when he or she is recumbent. In many cases, much greater anatomical detail is shown in a standing scan, and causes of a patient's symptoms that would go undiagnosed if the patient is scanned in a lying position might be correctly identified if the same patient is scanned while standing. Thus, there may be great benefit to having a patient stand during the scanning procedure.
However, standing for the duration of a multiple-slice scan procedure is even more demanding than during an imaging session that is conducted while the patient is lying down. A standing patient is more likely to become fatigued, and to fidget or shift his or her weight during the course of the scan. Because the patient is likely being scanned to analyze a physical ailment, he or she would probably have difficulty in enduring a standing multiple-slice procedure. Thus, in order to take advantage of the benefits offered by scanning a standing patient, it is necessary to speed up the scanning procedure.
Conventional multiple-slice scan procedures have other drawbacks as well. For example, the distance between slice boundaries in the direction of the thickness of the slices must be limited, in order to avoid saturation. A result of closing the gap between slice boundaries is a drop in signal-to-noise ratio, which in turn leads to images having a resolution that is unsuitable for reliable medical diagnosis. A larger inter-slice boundary, however, also lessens the utility of the slice images and the dependability of the resulting analysis.
Accordingly, a significant need exists for shortening the time period for obtaining the T1-weighted and T2-weighted images, and in particular, to reduce the total acquisition time for acquiring the data and information from which T1 weighted and T2-weighted NMR images are constructed. Further, there is a need for an overall scanning methodology that can provide for more closely-spaced slice acquisition data.
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