fig1 shows , diagrammatically a magnetic resonance device 1 according to the invention comprising a magnetic field generator 2 for generating a stationary homogeneous magnetic field , an rf radiator 3 for radiating radio frequency pulses to an object 4 , a modulator 5 for modulating the rf - pulses , a receiver and demodulator 6 for receiving and demodulating excited magnetic resonance signals comprising a sampler 7 for sampling the magnetic resonance signals , and a controller 8 for controlling said rf radiator , modulator , receiver / demodulator and sampler 3 , 5 , 6 and 7 . the device 1 further comprises a processor 9 for processing the sampled magnetic resonance signals , and a gradient generator 10 for generating magnetic field gradients which are superimposed upon the stationary magnetic field . in further detail , the radiator 3 comprises the modulator 5 , a power amplifier 11 , a directional coupler and a transmitter and receiver coil 13 . the modulator 5 comprises a modulator 14 and an oscillator 15 . the modulation can be amplitude , frequency or phase modulation or a combination thereof , controlled by the controller 8 . the radiator 3 excites nuclear spins in the object 4 when transmitting rf - pulses , and the receiver and demodulator 6 will receive magnetic resonance signals in response thereto . the transmission and reception may have various embodiments , e . g . the coil 13 may be a transmitter / receiver coil or may be embodied as separate transmitter and receiver coils . in the separate embodiment no directional coupler 12 will be present . the receiver and demodulator 6 may comprise a quadrature demodulator , in which case the sampler 7 will comprise two analog - to - digital convertors supplying quadrature detected samples of respective sin and cos components of the magnetic resonance signal . the gradient generator 10 comprises a power supply 16 controlled by the controller 8 , for separate control of gradient magnetic coils 17 , 18 and 19 which generate magnetic field gradients g x , g y , and g z respectively . the gradient coils 17 , 18 and 19 may be arranged such that the field direction of the magnetic field gradients coincides with the direction of the stationary homogeneous magnetic field , and that the directions of the gradients are perpendicular to each other , indicated in fig1 with mutually perpendicular axes x , y and z . if the magnetic resonance device 1 is put into operation and the object 4 is placed inside the means 2 for generating the stationary field , a small excess of nuclear spins ( of nuclei exhibiting a magnetic moment ) will be directed in the direction of the stationary magnetic field , which can be considered macroscopically as a magnetization m 0 , an equilibrium magnetization . the equilibrium magnetization m 0 can be disturbed by the application of pulse and gradient sequences to the object 4 . for obtaining physiological information from the object 4 , the magnetic resonance device 1 comprises several sensors , such as sensors c , for picking up cardiac signals which are fed to preamplifiers 20 and 21 . the preamplified cardiac signals are fed to an ecg trigger unit 22 via a fibre optic cable 23 by means of an ecg transmitter 24 . the ecg trigger unit 22 supplies trigger pulses tr to the controller 8 for triggering pulse and gradient sequences . the ecg trigger unit 22 is further controlled by the controller 8 by means of control signal ctl for inhibiting the ecg trigger unit 22 from sending pulses during predetermined periods . the processor 9 may comprise a programmer so that the pulse and gradient sequence starts from a predetermined time from the trigger pulse tr . the trigger pulse tr may be derived in the trigger unit 22 from the r - wave of the electrocardiagram obtained from the object 4 . for a more detailed description of a mri device and a general description of the principles of mri , reference is made to the handbook , &# 34 ; practical nmr imaging &# 34 ;, m . a . foster and j . m . s . hutchinson , 1987 irl press . in fig2 a , a known multi heart phase myocardial tagging sequence tsq is shown as a function of the time t . at t = t 0 a tagging section tag is applied , e . g . a spamm tagging section as described in the cited axel article . the tagging section is applied under the control of the controller 8 a predetermined time interval after the controller receives a trigger pulse tr from the trigger means 22 . for obtaining a stripe - like tagging pattern , the tagging section comprises two radio frequency pulses separated by a &# 34 ; wrap &# 34 ; gradient . for obtaining a grid - like tagging pattern , the tagging section may comprise two such rf - pulse gradient combinations in succession . the two one dimensional tagging sections must be separated by so called spoiler or dephasing gradient , to obtain a multiplication of these two tagging patterns . the tagging section tag is followed by a multi heart phase imaging section after a time interval δ t , the imaging section comprising a number of gradient echo imaging segments temporally spaced over the cardiac cycle . each imaging segment comprises an rf - pulse p1 with excitation angle α , a slice selection gradient g1 , a phase encoding gradient g2 and a readout gradient g3 applied by the rf - coil 13 and the gradient coils 17 , 18 and 19 , respectively . the pulses α k , k = 0 ,..., n - 1 ( n : total number of heart phase images ), are low flip angle pulses , not necessarily all equal . the respective pulse angles can be optimized for particular repetition times and tissue longitudinal relaxation times . the sequence tsq is repeated a number of times , e . g . 256 , for different values of the phase encoding gradient g2 . the received and sampled magnetic resonance signals around t = t k , t = t k + 1 , ... are processed by the processor 9 , a tagged image being obtained for each heart phase . the time interval between the kth and ( k + 1 ) th imaging segment is δ k . in fig2 b , a spamm tagging section tag for applying a grid like tagging pattern is shown , comprising first and second radio frequency pulses p2 and p3 separated by a first &# 34 ; wrap &# 34 ; gradient g7 , and comprising third and fourth radio frequency pulses p4 and p5 separated by a second &# 34 ; wrap &# 34 ; gradient g6 , the first and the second &# 34 ; wrap &# 34 ; gradient being orthogonal with respect to each other . the gradients g4 and g5 , which must be in slice selection direction , are applied to dephase any x and y components of the tagging magnetization . the factors β and γ are determined by the integrals of the gradients g7 and g6 , respectively . the amplitudes of the pulses p2 , p3 , p4 and p5 can be constant . the radio frequency pulses p2 and p3 , and p4 and p5 form so - called &# 34 ; 11 &# 34 ; sets of pulses then , the sum of the pulse angles of the pulses p2 and p3 , and p3 and p4 being in a range of 90 ° and 180 °. binomial pulse combinations such as &# 34 ; 121 &# 34 ; or &# 34 ; 1331 &# 34 ; pulses can also be applied , separated by gradients . practically , the summed angle amounts between 90 ° and 180 °. for a specified tissue , a specified number of excitation pulses of a specified angle and for a specified time point , the tagging angle can be optimized to obtain a better contrast than a known spamm sequence . for example , when measuring a diastolic , a midstolic , and a systolic modulus image with an excitation angle of 35 °, the tagging contrast at the systolic images on the myocardium is maximal , using a tagging angle of about 140 °. the present invention overcomes such problems . application of binomial pulse combinations might take too long , i . e . the tagging section might not be completed before the heart begins to move . before a more detailed description of the present invention will be given , shortcomings of the known spamm tagging will first be outlined . it is assumed for simplicity that the object 4 to be imaged has a uniform magnetization density equal to unity . then , immediately after the tag section , the z - magnetization is given by the shape of each heart phase image is determined by the z magnetization immediately before the respective rf excitation pulses p1 . for the first image this z - magnetization is represents a decaying image of the object which has been modulated by the tagging pattern , and the second term , represents an image of the object due only to that magnetization that has been relaxed since the application of the tagging section tag . for the kth heart phase image qt k contains the information for monitoring myocardial motion , while qr k is a term i which reduces the contrast of the tagging grid , and for a tagging section with a 90 ° tagging angle it can easily be shown that λk being dependent on the αj &# 39 ; s , δj &# 39 ; s , t 1 etc ., but not depending on the spatial coordinates x and y . in the presence of motion the variables x and y represent object coordinates at the point in time t = t 0 at which the tagging section tag is applied . the tagged image is given by basically the tagging grid can be automatically detected by searching for the lines of minima within the tagged image . however , due to the presence of the term qr k and the influence of λk , the actual values of these minima vary widely . in an image with large intensity variations , which is usually the case , it could be very difficult to detect these minima . and for modulus type images , as is dealt with in many cases , the image shape is as follows , in the latter case it does not suffice to search for lines of minima , their position being influenced by λk and qr k . the tagging grid , which is positioned between the lines of minima , can only be assigned definitely in homogeneous regions of the object in this case . according to the present invention , the term qr k , is eliminated in order to allow for reliable automatic detection of tagging grids in tagged images . the elimination is achieved by performing pairs of measurements with two spatially differing tagging functions and by subtracting the tagged images resulting from two tagging pulse and gradient sequences tsq with different tagging sections tag incorporating these tagging functions . denoting the respective tagging functions by tag1 ( x , y ) and tag2 ( x , y ), the best results are obtained when for a 180 ° spamm tagging section tag as shown in fig2 b this tagging pattern can inverted by negating the phase of any one of the four radio frequency pulses p2 , p3 , p4 or p5 , i . e . by replacing one of the four 90 ° pulses in &# 34 ; 11 &# 34 ; sets of pulses p2 , p3 and p4 , p5 respectively by a - 90 ° pulse . with this choice of tagging functions , the form of the tagged image is as follows by addition of the two tagged images resulting from the two tagging sequences an image with t 1 dependent anatomical contrast will be obtained with the form of according to the present invention , the lines of minima substantially have zero intensity , which make them very easy to detect . the zero intensity lines are detected automatically by means of the processor 9 through scanning of the resulting tagged image . furthermore , the tagging magnetic resonance method according to the present invention gives very sharp minima when the tagging grid is defined as a modulus image . in case of 90 ° tagging angles , according to the present invention , the tagging grid lines can be narrowed by applying binomial pulses , with a summed pulse angle of 90 ° for each tagging direction . the form of the tagging function then is , by rough approximation n being the order of the binomial sequence . the duration of the tagging sequence increases with the order n and tagging might not be completed before the heart starts to move . in case of myocardial motion the 180 ° variant is thus preferred .