Magnetic resonance method of measuring kidney filtration rates

A magnetic resonance imaging system employs a sequence of radio frequency pulses and magnetic field gradients to detect and measure the spin-relaxation time of moving blood within a subject. Spin-lattice relaxation times are determined by first inverting longitudinal spin magnetization and then detecting the recovery of this magnetization with a series of detection radio frequency pulses. The inversion pulse is applied to the entire subject, but the detection pulses are applied only to a selection portion of the subject. Blood motion causes the blood in the detection region to be replaced for each detection pulse, thereby increasing the accuracy of the measurement and permitting the use of multiple detection pulses after a single inversion pulse. In-vivo application of this invention can be used to assess renal function in individual kidneys.

CROSS REFERENCE TO RELATED APPLICATIONS 
This application is related to applications "Methods for measurement of 
Longitudinal Spin Relaxation Times in Moving Liquids" (Ser. No. 
08/185,256), "Methods for High Speed Measurement of Spin-Lattice 
Relaxation Times" (Ser. No. 08/105,249) both by Charles L. Dumoulin, filed 
concurrently with this application, and assigned to the present assignee. 
BACKGROUND OF THE INVENTION 
1. Field of the Invention 
The present invention relates to the fields of nuclear magnetic resonance 
spectroscopy and magnetic resonance (MR) imaging, and more specifically 
longitudinal spin relaxation time measurement, and the in-vivo assessment 
of renal function. 
2. Discussion of Prior Art 
Presently, the measurement of in-vivo spin-lattice relaxation times is a 
useful procedure in diagnostic radiology. In these procedures a subject is 
placed in a magnet to generate longitudinal spin magnetization in 
resonating nuclei of the subject or "nuclear spins". In the most commonly 
used procedure (known as Inversion Recovery) this magnetization is 
inverted by the application of a radio frequency pulse capable of nutating 
the longitudinal spin magnetization 180.degree. . When the magnetization 
of the subject's nuclear spins is inverted, it spontaneously returns to 
the non-inverted equilibrium state. The return to the equilibrium state 
occurs in an exponential fashion having a half-life which is 
characteristic of the molecular environment of the nuclear spin. This 
half-life is frequently given the name longitudinal spin relaxation time, 
T.sub.1. 
During the return to the equilibrium (or fully relaxed) state, the 
longitudinal magnetization cannot be directly detected. The instantaneous 
mount of longitudinal magnetization can be measured, however, by applying 
a sampling RF pulse. This sampling RF pulse nutates the longitudinal 
magnetization into the transverse plane, thereby creating transverse spin 
magnetization. Maximum transverse spin magnetization is generated by the 
application of a 90.degree. nutation. Unlike longitudinal magnetization, 
transverse spin magnetization is capable of inducing a signal in a 
receiver coil placed near the subject. Once transverse spin magnetization 
is generated, it can be phase shifted using magnetic field gradient pulses 
of selected intensities and durations. These gradient-induced phase shifts 
encode the position of spin magnetization within the magnet. Two or 
three-dimensional images of the distribution of spin magnetization can be 
generated by repeating the sequence of RF and magnetic field gradient 
pulses and acquiring the MR signal responsive to a collection of magnetic 
field gradient intensities. 
In-vivo measurement of T.sub.1 with previously available methods typically 
requires a long acquisition time. This is because the longitudinal 
magnetization must be measured at multiple points in time after the 
inversion pulse to accurately determine the half-life of the recovery. 
Only a single sampling pulse can be used during the recovery process. This 
is because application of a sampling pulse disturbs the longitudinal spin 
magnetization, and thus compromises the integrity of measurements 
generated by any subsequent sampling pulses. Furthermore, best results are 
obtained when full recovery of longitudinal spin magnetization occurs 
after each sampling pulse. For in-vivo applications the time for full 
relaxation is between 1500 and 5000 ms, since most in-vivo T.sub.1 values 
are between 300 and 1000 ms. Measurement of T.sub.1 for each pixel in an 
image may require exam times as long as an hour, since enough data must be 
acquired to construct an image (typically with a resolution of 
256.times.256), for each of several sampling times (typically 4-8) after 
each inversion pulse. 
An alternative method for in-vivo T.sub.1 measurement described by Campeau 
et. al. in the Proceedings of the Eleventh Annual Meeting of the Society 
of Magnetic Resonance in Medicine, 1992, pg. 434, employs a series of 
slice selective inversion pulses which excite slices placed orthogonal to 
the image plane of an acquired MR image. Each inverted slice is in a 
unique location and each inversion pulse is applied at a unique time 
before the application of the transverse spin magnetization generation 
pulse of the imaging pulse sequence. If the acquired image has relatively 
large features of homogeneous T.sub.1, (e.g. a large skeletal muscle) the 
resulting image will contain a series of stripes, each created by spin 
inversion at selected times prior to the application of the detection 
pulse. The T.sub.1 values of the selected image feature can then be 
determined by measuring the pixel intensity in each stripe corresponding 
to each inversion delay time and fitting the result to an exponential 
equation to determine the rate constant, T.sub.1. While this method is 
relatively fast, it is not suited for the T.sub.1 measurement of small 
features such as the blood in a selected blood vessel. The technique is 
also poorly suited for T.sub.1 measurement of moving blood, since blood 
motion during the period between each selective inversion pulse and the 
detection pulse causes mixing of the inverted boluses of blood. 
Present clinical techniques used to assess renal function are based on the 
concept of clearance. Under normal steady-state conditions, the daily 
production of creatinine is equal to its daily excretion, thereby 
regulating serum creatinine within a narrow range. Thus, clinicians 
frequently use serum creatinine concentration alone as an estimate of the 
Glomerular Filtration Rate (GFR). This technique, however, has limited 
accuracy and the presence of unilateral kidney disease is usually not 
detectable. Since this test may detect a normal serum creatinine 
concentration even in the presence of a 50% reduction in GFR, renal 
insufficiency may be misdiagnosed. 
Currently there is a need for a non-invasive method of measuring the 
longitudinal spin relaxation time of moving liquids for use in application 
such as in assessing renal function. 
OBJECTS OF THE INVENTION 
An object of the present invention is to provide a system which is capable 
of rapid longitudinal spin relaxation time, T.sub.1, measurements. 
Another object of the present invention is to provide a system which 
measures the T.sub.1 of blood in selected blood vessels. 
Another object of the present invention is to provide a method in which 
blood velocity measurements are made of a volume of moving blood 
simultaneously with T.sub.1 measurements of the same blood volume. 
Another object of the present invention is to provide a method in which 
T.sub.1 measurements of moving blood are employed in assessing renal 
function. 
SUMMARY OF THE INVENTION 
In the present invention, a subject is placed in a Magnetic Resonance 
Imaging (MRI) system. A novel inversion recovery MR pulse sequence is then 
used to measure the T.sub.1 of moving blood. This pulse sequence employs 
an inversion pulse which is not spatially selective to invert all nuclear 
spins within a selected portion of the subject. The inversion pulse is 
followed by a series of detection pulses which nutate the longitudinal 
magnetization by as much as 90 degrees. These detection pulses are 
spatially selective and in the present invention have a slice profile. 
Each detection pulse is applied in the same location in the subject. This 
location is chosen to transect a selected blood vessel in which blood is 
flowing. Since all the blood in the subject exhibits inverted longitudinal 
spin magnetization and since blood motion replaces blood in the detection 
slice after each detection pulse, longitudinal spin magnetization is 
accurately sampled. In addition, sampling of longitudinal spin 
magnetization is performed a plurality, N, times during the recovery 
process, instead of once as in prior methods. This results in an N-fold 
reduction in the total scan time. 
In one embodiment of the present invention velocity-encoding magnetic field 
gradient pulses are incorporated into the pulse sequence to permit the 
measurement of blood velocity simultaneously with the measurement of blood 
T.sub.1. 
The novel pulse sequence of the present invention can be used for the 
assessment of single kidney hemodynamic function. For example, after 
injection of an MR contrast agent such as Gadolinium diethylene triamine 
pentaacetic acid (Gd-DTPA) into the subject, the agent is excreted through 
the subject's kidneys. The efficiency of renal filtration of the contrast 
agent, or Filtration Fraction (FF), can be derived by measuring the 
T.sub.1 of blood in the arterial system (before it passes through the 
kidney) and the renal vein. When the filtration fraction measurement is 
combined with blood velocity measurements in the renal vessels, the 
Glomerular Filtration Rate (GFR) can be determined.

DETAILED DESCRIPTION OF THE INVENTION 
In the present embodiment of the invention, a subject is placed within the 
magnet of a magnetic resonance imaging system. The region over which 
longitudinal spin relaxation time, T.sub.1, measurement is desired is then 
identified by an operator, perhaps with the assistance of a conventional 
MR imaging sequence. A pulse sequence is then applied and the data 
analyzed. 
FIG. 1 is a simplified block diagram of the major components of a magnetic 
resonance (MR) imaging system suitable for use with the invention 
described herein. The system is made up of a general purpose mini-computer 
2 which is functionally coupled to a disk storage unit 2a and an interface 
unit 2b. A radiofrequency (RF) transmitter 3, signal averager 4, and 
gradient power supplies 5a, 5b and 5c, are all coupled to computer 2 
through interface unit 2b. Gradient power supplies 5a, 5b, 5c energize 
gradient coils 12-1, 12-2, 12-3 to create magnetic field gradients 
G.sub.x, G.sub.y, G.sub.z, respectively, in the "X", "Y", "Z" directions, 
respectively, over a subject to be imaged. RF transmitter 3 is gated with 
pulse envelopes from computer 2 to generate RF pulses having the required 
modulation to excite an MR response signal from a subject. The RF pules 
are amplified in an RF power amplifier 6 to levels varying from 100 watts 
to several kilowatts, depending on the imaging method, and applied to a 
transmitter coil 14-1. The higher power levels are necessary for large 
sample volumes, such as in whole body imaging, and where short duration 
pulses are required to excite larger NMR frequency bandwidths. 
The MR response signal is sensed by a receiver coil 14-2, amplified in a 
low noise preamplifier 9 and passed to receiver 10 for further 
amplification, detection, and filtering. The signal is then digitized for 
averaging by signal averager 4 and for processing by computer 2. 
Preamplifier 9 and receiver 10 are protected from the RF pulses during 
transmission by active gating or by passive filtering. 
Computer 2 provides gating and envelope modulation for the MR pulses, 
blanking for the preamplifier and RF power amplifier, and voltage 
waveforms for the gradient power supplies. The computer also performs data 
processing such as Fourier transformation, image reconstruction, data 
filtering, imaging display, and storage functions (all of which are 
conventional and outside the scope of the present invention). 
Transmitter coil 14-1 and receiver RF coil 14-2, if desired, may comprise a 
single coil. Alternatively, two separate coils that are electrically 
orthogonal may be used. The latter configuration has the advantage of 
reduced RF pulse breakthrough into the receiver during pulse transmission. 
In both cases, the coils are orthogonal to the direction of a static 
magnetic field B.sub.0 produced by a magnet means 11. The coils may be 
isolated from the remainder of the system by enclosure in an RF shielded 
cage. 
Magnetic field gradient coils 12-1, 12-2, and 12-3 are necessary to provide 
gradients G.sub.x, G.sub.y, and G.sub.z, respectively, that are monotonic 
and linear over the sample volume. Multivalued gradient fields cause a 
degradation in the MR response signal data, known as aliasing, which leads 
to severe image artifacts. Nonlinear gradients cause geometric distortions 
of the image. 
Magnet assembly 11, shown schematically in FIG. 2, has a central 
cylindrical bore 11a which generates a static magnetic field B.sub.0, 
typically in the axial, or Z Cartesian coordinate direction. A set of 
coils 12, such as coils 12-1,12-2, 12-3 of FIG. 1, receive electrical 
signals via input connections 12a, and provide at least one gradient 
magnetic field within the volume of bore 11a. Also situated within bore 
11a is an RF coil 14, which receives RF energy via at least one input 
cable 14a, to provide an RF magnetic field b.sub.1, typically in the X-Y 
plane. 
FIG. 3 is a pulse sequence diagram of radio frequency (RF) pulses and 
magnetic field gradients employed in a first embodiment of the present 
invention which may be executed by the MR imaging system of FIGS. 1 and 2. 
Pulse sequence 100 consists of a non-selective inversion radio frequency 
(RF) pulse 110 which substantially inverts all spin magnetization within 
the excitation radio frequency coil of the imaging system. In the present 
embodiment of the invention inversion RF pulse 110 is applied without the 
simultaneous application of a magnetic field gradient pulse. If it is 
desirable to restrict the volume of inversion, a magnetic field gradient 
pulse can be applied simultaneously with an inversion pulse in a manner 
well known to those skilled in the art. 
At a selected time after inversion RF pulse 110 is applied, a first 
subsequence 120a is applied. Subsequence 120a is comprised of a detection 
RF pulse 130 which is applied in the presence of a slice selective 
magnetic field gradient pulse 140. Detection pulse 130 nutates spin 
magnetization in a selected portion of the subject. The amount of nutation 
can be selected by selecting the duration and amplitude of detection pulse 
130. Maximum transverse magnetization is most often obtained with a 
nutation angle of 90 degrees. The location and size of the selected 
portion can be adjusted by appropriate selection of the frequency and 
bandwidth of RF pulse 130 and the amplitude of slice selective magnetic 
field gradient pulse 140. In the present embodiment of the invention the 
selected portion has a planar geometry of selected thickness. In 
alternative embodiments, a non-planar geometry such as a cylinder, can be 
used as described in U.S. Pat. No. 5,133,357 "Quantitative Measurement of 
Blood Flow Using Cylindrically Localized Fourier Velocity Encoding" issued 
Jul. 28, 1992, assigned to the present assignee, and hereby incorporated 
by reference. 
After the detection RF pulse 130 and slice selective magnetic field 
gradient pulse 140 are applied, a slice refocusing magnetic field gradient 
pulse 150 is applied. Slice refocusing gradient pulse 150 has an amplitude 
and duration which is selected to cause all transverse spin magnetization 
within the selected portion of the subject to be substantially in phase 
after the application of slice refocusing gradient pulse 140. In the 
present embodiment the product of the amplitude and duration of slice 
refocusing gradient pulse 150 is substantially half that of the negative 
of the product of the amplitude and duration of slice selective gradient 
pulse 140 in a manner well known to those skilled in the art. 
After detection RF pulse 130 and slice selective gradient pulse 140 have 
been applied, a phase encoding magnetic field gradient pulse 160 of a 
selected amplitude is applied. Phase encoding gradient pulse 160 is 
applied in a direction substantially orthogonal to slice selective 
gradient pulse 140 and can be applied simultaneously with slice refocusing 
pulse 150 if desired. For the sake of clarity, phase encoding pulse 160 
and slice refocusing pulse 150 are not shown to be simultaneous in FIG. 3, 
but it is possible to apply both simultaneously. 
After detection RF pulse 130 and slice selective gradient pulse 140 have 
been applied, a readout dephasing magnetic field gradient pulse 170 of a 
selected amplitude is applied. Readout dephasing gradient pulse 170 is 
applied in a direction substantially orthogonal to both slice selective 
gradient pulse 140 and phase encoding pulse 160. Readout dephasing pulse 
170 can be applied simultaneously with either slice refocusing pulse 150 
or phase encoding pulse 160 if desired. Readout dephasing pulse 170 causes 
transverse magnetization at different positions along the direction of the 
readout dephasing magnetic field gradient to obtain phase shifts which are 
proportional to position in the readout direction. 
Following the application of slice refocusing pulse 150, phase encoding 
pulse 160 and readout dephasing pulse 170, a readout magnetic field 
gradient pulse 180 is applied. Readout pulse 180 is applied in the same 
direction as readout dephasing pulse 170, but is given the opposite 
polarity. The amplitude and duration of readout pulse 180 is selected so 
that substantially all transverse spin magnetization has an identical 
phase shift at a selected point during readout pulse 180. 
Substantially simultaneously with readout pulse 180, a data acquire signal 
pulse 190 is sent to a data acquisition subsystem which is part of the 
imaging system. MR signals are digitized during data acquire pulse 190. 
Since the MR signals coming from transverse spin magnetization within the 
selected portion of the subject are acquired during readout magnetic field 
gradient 180, each detected MR signal will have a frequency which is 
proportional to the location of the position of the transverse spin 
magnetization which generated said signal. The location of each signal 
source can be determined by applying a Fourier transformation to the 
acquired signal data in a fashion well known to those skilled in the art. 
At a selected time interval after the application of subsequence 120a, a 
second subsequence 120b is applied. Subsequence 120b is substantially 
identical to subsequence 120a. Data acquired responsive to subsequence 
120b, however, is stored in its own location. Subsequence 120b is followed 
after a selected interval by a third subsequence 120c and so on until a 
plurality, N, subsequences have been applied. In the embodiment of the 
invention illustrated in FIG. 3 the interval between each subsequences is 
the same. In alternative embodiments the intervals can be arbitrarily 
chosen by the operator. 
Pulse sequence 100 is repeated a plurality, Y, times. In each repetition of 
pulse sequence 100, phase encoding pulse 160 is given a different 
amplitude. For each repetition of pulse sequence 100, however, the 
amplitude of phase encoding pulse 160 is identical for each subsequence 
120a-120N. Phase encoding pulse 160 causes phase shifts in the detected MR 
signals which are proportional to the position of transverse spin 
magnetization along the direction of phase encoding magnetic field 
gradient 160. Data acquired responsive to different amplitudes of phase 
encoding gradient 160 can be Fourier transformed to give the position (in 
the direction of phase encoding gradient 160) of the signal producing 
transverse spin magnetization in a manner well known to those skilled in 
the art. 
Once data has been collected responsive to Y repetitions of N subsequences, 
a total of N magnetic resonance images can be constructed. These N images 
will have pixels whose intensifies vary as a function of a time interval, 
t, between inversion RF pulse 110 and detection RF pulse 130 of each 
respective subsequence. The signal intensity, I, for a pixel having a 
single recovery rate, T.sub.1, can be described by the equation: 
EQU I=A+B*exp(-t/T.sub.1) [1] 
where A and B are constants. T.sub.1, A and B of any desired pixel or 
collection of pixels can be determined by fitting equation 1 to the 
intensities of the pixel obtained in each of the subsequences. This can be 
done using a non-linear least square optimization procedure such as that 
described in "Numerical Recipes in FORTRAN" by W. H. Press, S. A. 
Teukolsky, W. T. Vetterling and B. P. Flannery, pp. 678-683. 
Note that for portions of the subject which are relatively stationary, 
detection RF pulse 130 is repeatedly applied after inversion RF pulse 110. 
This causes the intensity of the detected MR signal for stationary tissue 
to have complicated characteristics which prevent an accurate analysis of 
T.sub.1. Blood contained within arteries and veins of the subject is in 
motion, however, and as it passes through the detection region, it is 
subjected only to non-selective inversion RF pulse 110 and a single 
detection RF pulse 130. For this to happen the blood must be moving at a 
velocity, V, which is greater than D/T where D is the thickness of the 
detection region and T is the time interval between detection pulses. 
FIG. 4 illustrates a second embodiment of the present invention. Like the 
embodiment described above in FIG. 3, pulse sequence 200 of FIG. 4 has an 
inversion RF pulse 210 followed by a plurality, N, of subsequences 
220a-200N, Each pulse sequence is further comprised of a detection RF 
pulse 230, a slice select magnetic field gradient pulse 240, slice 
refocusing magnetic field gradient pulse 250, a phase encoding magnetic 
field gradient pulse 260, a readout dephasing magnetic field gradient 
pulse 270, a readout magnetic field gradient pulse 280 and a data acquire 
signal pulse 290. 
In addition to the RF and magnetic field gradient pulses found in the first 
embodiment of the present invention, each subsequence 220a-220N in the 
second embodiment also has a two-lobed flow-encoding magnetic field 
gradient pulse. The flow-encoding pulse consists of a first flow-encoding 
magnetic field gradient pulse lobe 255a and a second flow-encoding 
magnetic field gradient pulse lobe 255b. The product of the pulse duration 
and amplitude of second flow-encoding pulse lobe 255b is substantially 
equal to the negative of the product of the pulse duration and amplitude 
of first flow-encoding pulse lobe 255a. 
Successive application of first flow-encoding pulse lobe 255a and second 
flow-encoding pulse lobe 255b to transverse spin magnetization causes a 
phase shift in the magnetization which is proportional to the velocity 
component of the magnetization parallel to the direction of the 
flow-encoding magnetic field gradient. This phase shift can be used to 
distinguish moving from stationary transverse spin magnetization. The 
phase of each portion of transverse spin magnetization, however, will have 
contributions from sources other than velocity. These sources may include 
transmitter offsets, chemical shift effects and eddy currents. 
In order to remove contributions from all components other than velocity, 
pulse sequence 200 is repeated and a second data set acquired. The RF and 
magnetic field gradient pulses of the repeated pulse sequence are 
identical to that of the first with the exception of first flow-encoding 
pulse lobe 255a and second flow-encoding pulse lobe 255b. In their place a 
third flow-encoding pulse lobe 255c followed by a fourth flow-encoding 
pulse lobe 255d applied. Third and fourth flow-encoding pulse lobes 255c, 
255d are identical to first and second flow-encoding 255a, 255b 
respectively, except that they have opposite polarity. Phase shifts 
induced by the third and fourth flow-encoding gradient lobes will also 
have opposite polarity relative to each other. Consequently, when the 
phase of data acquired responsive to the first application of pulse 
sequence 200, is subtracted from the phase of the data acquired responsive 
to the repeated application of pulse sequence 200, phase contributions 
from all non-velocity sources are substantially canceled, leaving only a 
phase shift arising from velocity. This phase shift is directly 
proportional to velocity and can be used to quantify velocity. Note that 
MR images generated with pulse sequence 200 have velocity information 
encoded as phase and T.sub.1 information encoded as magnitude. 
The present invention can be used to evaluate renal function in living 
systems as illustrated in FIG. 5. In FIG. 5, a kidney 300 present within a 
living subject is perfused with arterial blood 320 which flows into a 
renal artery 300a. Kidney 300 removes substances from the blood, and 
returns venous blood 330 to the body through a renal vein 300b. Substances 
which are removed from arterial blood 320 are passed through a ureter 300c 
to the bladder as urine 340. The present invention can be used to assess 
renal efficiency by measuring the concentration of substances within the 
subject's bloodstream before it enters, and after it exits the kidney, 
being prefiltered, and post filtered blood. This is easily done by 
selection of a suitable slice selection magnetic field gradient 140 and 
detection RF pulse 130 of FIG. 3 to form an image plane 310. If desired, 
the present invention can be performed with multiple image planes to 
concurrently obtain diagnostic information from more than a single 
location. This may be particularly useful in making independent 
assessments of renal function in two kidneys. 
To evaluate glomerular filtration rates, both the filtering efficiency of 
the kidney and the blood flow through the kidney must be determined. The 
filtration fraction, FF, can be determined with the present invention by 
introducing a longitudinal spin relaxation contrast agent into the 
bloodstream of the subject. Filtration Fraction can be expressed as: 
##EQU1## 
where C.sub.a is the concentration of the T.sub.1 relaxation contrast 
agent in the arterial blood entering the kidney and C.sub.v is the 
concentration of the relaxation contrast agent in the venous blood as it 
leaves the kidney. 
Gadolinium diethylene triamine pentaacetic acid (Gd-DTPA) is a commonly 
used relaxation contrast agent which is excreted primarily by glomerular 
filtration in the kidneys. Gd-DTPA, and other T.sub.1 relaxation contrast 
agents, shorten the measured T.sub.1 of blood by an amount which is 
substantially inversely proportional to the concentration of the T.sub.1 
relaxation contrast agent as illustrated in FIG. 6. In FIG. 6, the T.sub.1 
of blood in the absence of any T.sub.1 relaxation contrast agent (i.e. 
concentration =0) is denoted as T.sub.1B. At a concentration, C, the 
T.sub.1 relaxation contrast agent reduces the T.sub.1 value of blood to 
T.sub.1C. Blood having a reference concentration of T.sub.1 relaxation 
contrast agent, R, will have a relaxation time of T.sub.1R. The observed 
relaxation time of blood, T.sub.1 (observed) for a given concentration, C, 
of T.sub.1 relaxation contrast agent can be expresses as: 
##EQU2## 
This equation can be rearranged to give: 
##EQU3## 
Equation [4] can be combined with equation [2] if C.sub.A is derived from 
the observed T.sub.1 of the arterial blood and C.sub.V is derived from the 
observed T.sub.1 of the venous blood. The resulting expression for 
filtration fraction becomes: 
##EQU4## 
where T.sub.1V is the observed T.sub.1 of venous renal blood and T.sub.1A 
is the observed T.sub.1 of arterial renal blood. 
The accuracy of the measurement of filtration fraction depends on the 
accuracy of the T.sub.1 measurement and on the magnitude of the difference 
between T.sub.1A and T.sub.1B. Higher concentrations of T.sub.1 relaxation 
contrast agent in the arterial blood will result in a greater reduction in 
the T.sub.1 of blood thereby increasing the denominator of equation [5]. 
This effect manifests itself as an increase in the dynamic range of FF as 
illustrated in FIG. 7. 
FIG. 7 illustrates the relationship between T.sub.1A and T.sub.1V for 
Filtration Fractions of 0.0, 0.1, 0.2 and 0.3. Note that if the Filtration 
Fraction is zero, the T.sub.1 of blood passing into the kidney is 
identical to the T.sub.1 of blood exiting the kidney since no T.sub.1 
relaxation contrast agent is removed. For higher values of Filtration 
Fraction, the concentration of the T.sub.1 relaxation contrast agent in 
the venous renal blood is lower than that found in the arterial blood. 
Consequently, the T.sub.1 of venous blood is longer. 
Once the Filtration Fraction has been measured, the Glomerular Filtration 
Rate, rate.sub.GF, can be determined from the expression: 
EQU rate.sub.GF =FF*flow.sub.RB *(1-Hct) [6] 
where flow.sub.RB is the Renal Blood Flow and Hct is the hematocrit, or 
fraction of the blood which is comprised of red blood cells. 
FIG. 8 shows an alternative embodiment of the present invention in which a 
high resolution Nuclear Magnetic Resonance (NMR) system is used to obtain 
NMR spectra. Here a sample solution is circulated through a magnet 610 via 
connecting tube 630. A pump means 620 is used to move the sample solution 
through the detection plane created by slice selection magnetic field 
gradient pulse 140 and detection RF pulse 130 (FIG. 3). Phase-encoding 
gradient pulse 160, readout dephasing gradient pulse 170 and readout 
gradient pulse 180 (FIG. 3) can be omitted to permit the detection of an 
NMR spectrum rather than an MR image. Alternatively, other pulse sequences 
known in the art of NMR spectroscopy can be employed if desired. 
Reduction to Practice 
The present invention has been reduced to practice. FIG. 9 is a graph of a 
T.sub.1 recovery curve of signal intensity measured in the renal artery of 
a healthy human subject prior to the injection of a T.sub.1 relaxation 
contrast agent. FIG. 10 is a graph of a T.sub.1 recovery curve of signal 
intensity measured in the renal artery after the injection of a T.sub.1 
relaxation contrast agent (Gd-DTPA). FIG. 11 is a graph of a T.sub.1 
recovery curve of signal intensity measured in the renal vein after the 
injection of the T.sub.1 relaxation contrast agent. The T.sub.1 of blood 
prior to the T.sub.1 relaxation contrast agent injection was measured to 
be 1229 ms. After injection, the T.sub.1 of arterial blood was found to be 
568 ms while the T.sub.1 of venous blood was found to be 823 ms. These 
relaxation times were entered into equation [5] to obtain a filtration 
fraction equal to 0.26. This filtration fraction is well within the 
expected value for a healthy individual. 
While several presently preferred embodiments of the novel T.sub.1 
measurement system have been described in detail herein, many 
modifications and variations will now become apparent to those skilled in 
the art. It is, therefore, to be understood that the appended claims are 
intended to cover all such modifications and variations as far within the 
true spirit of the invention.