WO2014147589A1

WO2014147589A1 - Apparatus and method for influencing and/or detecting magnetic particles comprising compensation unit

Info

Publication number: WO2014147589A1
Application number: PCT/IB2014/060021
Authority: WO
Inventors: Ingo Schmale; Bernhard Gleich
Original assignee: Koninklijke Philips N.V.; Philips Deutschland Gmbh
Priority date: 2013-03-21
Filing date: 2014-03-21
Publication date: 2014-09-25

Abstract

The present invention relates to an apparatus and a method for influencing and/or detecting magnetic particles in a field of view (28). The apparatus comprises selection device for generating a magnetic selection field (50) and drive and receiving device comprising one or more combined drive field and selection coils (330) for changing the position in space of the two sub-zones (52, 54) in the field of view (28) and for acquiring detection signals. A compensation unit (320) comprising a measurement inductor (321) and a gradiometer inductor (322) is provided, said measurement inductor (321) and said gradiometer inductor (322) being inductively coupled, wherein said compensation unit (320) is electrically coupled to said drive field signal generator unit (122), said signal receiving unit (140) and a drive-receiving coil (330).

Description

Apparatus and method for influencing and/or detecting magnetic particles comprising compensation unit

FIELD OF THE INVENTION

The present invention relates to an apparatus and a method for influencing and/or detecting magnetic particles in a field of view. Further, the present invention relates to a computer program for implementing said method on a computer and for controlling such an apparatus. The present invention relates particularly to the field of Magnetic Particle Imaging.

BACKGROUND OF THE INVENTION

Magnetic Particle Imaging (MPI) is an emerging medical imaging modality. The first versions of MPI were two-dimensional in that they produced two-dimensional images. Newer versions are three-dimensional (3D). A four-dimensional image of a non- static object can be created by combining a temporal sequence of 3D images to a movie, provided the object does not significantly change during the data acquisition for a single 3D image.

MPI is a reconstructive imaging method, like Computed Tomography (CT) or

Magnetic Resonance Imaging (MRI). Accordingly, an MP image of an object's volume of interest is generated in two steps. The first step, referred to as data acquisition, is performed using an MPI scanner. The MPI scanner has means to generate a static magnetic gradient field, called the "selection field", which has a (single or more) field-free point(s) (FFP(s)) or a field-free line (FFL) at the isocenter of the scanner. Moreover, this FFP (or the FFL; mentioning "FFP" in the following shall generally be understood as meaning FFP or FFL) is surrounded by a first sub-zone with a low magnetic field strength, which is in turn surrounded by a second sub-zone with a higher magnetic field strength. In addition, the scanner has means to generate a time-dependent, spatially nearly homogeneous magnetic field. Actually, this field is obtained by superposing a rapidly changing field with a small amplitude, called the "drive field", and a slowly varying field with a large amplitude, called the "focus field". By adding the time-dependent drive and focus fields to the static selection field, the FFP may be moved along a predetermined FFP trajectory throughout a "volume of scanning" surrounding the isocenter. The scanner also has an arrangement of one or more, e.g. three, receive coils and can record any voltages induced in these coils. For the data acquisition, the object to be imaged is placed in the scanner such that the object's volume of interest is enclosed by the scanner's field of view, which is a subset of the volume of scanning.

The object must contain magnetic nanoparticles or other magnetic non-linear materials; if the object is an animal or a patient, a tracer containing such particles is administered to the animal or patient prior to the scan. During the data acquisition, the MPI scanner moves the FFP along a deliberately chosen trajectory that traces out / covers the volume of scanning, or at least the field of view. The magnetic nanoparticles within the object experience a changing magnetic field and respond by changing their magnetization. The changing magnetization of the nanoparticles induces a time-dependent voltage in each of the receive coils. This voltage is sampled in a receiver associated with the receive coil. The samples output by the receivers are recorded and constitute the acquired data. The parameters that control the details of the data acquisition make up the "scan protocol".

In the second step of the image generation, referred to as image reconstruction, the image is computed, or reconstructed, from the data acquired in the first step. The image is a discrete 3D array of data that represents a sampled approximation to the position-dependent concentration of the magnetic nanoparticles in the field of view. The reconstruction is generally performed by a computer, which executes a suitable computer program. Computer and computer program realize a reconstruction algorithm. The reconstruction algorithm is based on a mathematical model of the data acquisition. As with all reconstructive imaging methods, this model can be formulated as an integral operator that acts on the acquired data; the reconstruction algorithm tries to undo, to the extent possible, the action of the model.

Such an MPI apparatus and method have the advantage that they can be used to examine arbitrary examination objects - e. g. human bodies - in a non-destructive manner and with a high spatial resolution, both close to the surface and remote from the surface of the examination object. Such an apparatus and method are generally known and have been first described in DE 101 51 778 Al and in Gleich, B. and Weizenecker, J. (2005), "Tomographic imaging using the nonlinear response of magnetic particles" in Nature, vol. 435, pp. 1214-1217, in which also the reconstruction principle is generally described. The apparatus and method for magnetic particle imaging (MPI) described in that publication take advantage of the non-linear magnetization curve of small magnetic particles.

MPI is based on the detection of harmonics as generated by magnetic (nano-) particles subjected to an external sinusoidal magnetic field excitation. Opposed to MR, excitation and reception are taking place simultaneously and are solely separated in the frequency domain. Conventionally, separation is realised by notch filters (LC resonators). Due to the higher sensitivity of coils that are nearest to the patient, there is a "competition" between drive (Tx-) and receive (Rx-) coil on the space very near around the patient.

SUMMARY OF THE INVENTION

It is an object of the present invention to provide an apparatus and a method for influencing and/or detecting magnetic particles in a field of view that achieves harmonic background reduction without reducing the size of the field of view, in particular the bore size of a bore of the apparatus into which a patient can be placed.

In a first aspect of the present invention an apparatus for influencing and/or detecting magnetic particles in a field of view is presented, which apparatus comprises:

selection device comprising a selection field signal generator unit and selection field elements for generating a magnetic selection field having a pattern in space of its magnetic field strength such that a first sub-zone having a low magnetic field strength where the magnetization of the magnetic particles is not saturated and a second sub-zone having a higher magnetic field strength where the magnetization of the magnetic particles is saturated are formed in the field of view,

drive and receiving device comprising a drive field signal generator unit, a signal receiving unit and one or more drive-receiving coils, a drive-receiving coil being configured both for changing the position in space of the two sub-zones in the field of view by means of a magnetic drive field so that the magnetization of the magnetic material changes locally and for acquiring detection signals, which detection signals depend on the magnetization in the field of view, which magnetization is influenced by the change in the position in space of the first and second sub-zone, and

a compensation unit comprising a measurement inductor and a gradiometer inductor, said measurement inductor and said gradiometer inductor being inductively coupled, wherein said compensation unit is electrically coupled to said drive field signal generator unit, said signal receiving unit and a drive-receiving coil.

In a further aspect of the present invention a corresponding method is presented.

In yet a further aspect of the present invention a computer program is presented comprising program code means for causing a computer to control an apparatus as according to the present invention to carry out the steps of the method proposed according to the present invention when said computer program is carried out on the computer.

Preferred embodiments of the invention are defined in the dependent claims. It shall be understood that the claimed method and the claimed computer program have similar and/or identical preferred embodiments as the claimed apparatus and as defined in the dependent claims.

Since there is crosstalk from the transmit side to the receive side, precisely harmonics generated from the power amplifier, a lot of effort is spent on the band pass filter as conventionally used to ensure that no harmonics from the drive field enter the receive path. However, it was found, that the success of this effort is limited finally by the component of the filter itself: particularly capacitors (but also other components and materials) behave non- linearly. The degree to which they are non-linear is so small that it is hard to measure and it seems to be of no concern to other applications. Nevertheless it becomes limiting for this reception scheme, despite efforts to identify an optimum capacitor technology.

One solution is to employ a gradiometer-based reception scheme, e.g. to have access to the fundamental frequency response of the magnetic particles. These gradiometer solutions are based on a dedicated receive coil, which receive the harmonic response from the magnetic particles (desired effect) but also the drive signal that excites these magnetic particles (undesired effect). To compensate the undesired drive signal, a second coil "balancing coil" is employed, which receives the same drive signal, but is connected with inverted polarity. For a bore-like drive-field generation, the dedicated receive coil is inside the transmit coils. Whilst this is beneficial with respect to signal strength (it has higher sensitivity being nearer to the magnetic particles), the strong disadvantage is that the bore size remaining for the patient (animal, object under investigation, ...) is reduced. The compensation coil needs to couple to the drive signal only. This can be achieved in two ways: either it is also within or near to the drive field generating transmit coils, or it couples to another external inductor through which the same current flows as through the drive field coil.

Thus, according to the present invention suppression of harmonic background in the receive path is achieved by using a gradiometric cancellation scheme, but without the reduction of bore size of the apparatus for placement of the patient by inset coils. For this purpose a compensation unit comprising a measurement inductor and a gradiometer inductor is used. The measurement inductor and a gradiometer inductor, both being preferably realized by a coil, are inductively coupled like an electrical transformer. Further, said compensation unit is electrically coupled between said drive field signal generator unit, said signal receiving unit and a drive-receiving coil (preferably all drive-receiving coils), said drive- receiving coil representing a combined drive field and receiving coil (i.e. functioning both as conventional drive field coil and as receiving coil).

Generally, according to the proposed magnetic particle imaging apparatus and method the magnetic gradient field (i.e. the magnetic selection field) is generated with a spatial distribution of the magnetic field strength such that the field of view comprises a first sub-area with lower magnetic field strength (e.g. the FFP), the lower magnetic field strength being adapted such that the magnetization of the magnetic particles located in the first sub- area is not saturated, and a second sub-area with a higher magnetic field strength, the higher magnetic field strength being adapted such that the magnetization of the magnetic particles located in the second sub-area is saturated. Due to the non-linearity of the magnetization characteristic curve of the magnetic particles the magnetization and thereby the magnetic field generated by the magnetic particles shows higher harmonics, which, for example, can be detected by a detection coil. The evaluated signals (the higher harmonics of the signals) contain information about the spatial distribution of the magnetic particles, which again can be used e.g. for medical imaging, for the visualization of the spatial distribution of the magnetic particles and/or for other applications.

Thus, the apparatus and the method according to the present invention are based on a new physical principle (i.e. the principle referred to as MPI) that is different from other known conventional medical imaging techniques, as for example nuclear magnetic resonance (NMR). In particular, this new MPI-principle, does, in contrast to NMR, not exploit the influence of the material on the magnetic resonance characteristics of protons, but rather directly detects the magnetization of the magnetic material by exploiting the non- linearity of the magnetization characteristic curve. In particular, the MPI-technique exploits the higher harmonics of the generated magnetic signals which result from the non-linearity of the magnetization characteristic curve in the area where the magnetization changes from the non- saturated to the saturated state.

According to a preferred embodiment said compensation unit is physically arranged separate from said drive-receiving coils. Thus, different from know apparatus, the measurement inductor and the gradiometer inductor are not arranged within or near the drive field coils (as provided in conventional apparatus) in which they consume much of the valuable space for placing the patient.

Preferably, said compensation unit is physically arranged at a distant location from said drive-receiving coils, e.g. as a completely separate unit arranged some meters away from the drive-receiving coils and/or outside a bore formed by said drive-receiving coils.

Further, in an embodiment said measurement inductor and said gradiometer inductor are wound as toroid. Such a toroid has the advantage of a small magnetic stray field to the outer environment and that several components can be arranged close together (e.g. three toroids, one for each channel, can be placed next to each other).

Still further, in an embodiment said compensation unit comprises, per drive- receiving coil, a measurement inductor and a gradiometer inductor being inductively coupled. Generally, all elements are provided at least once per channel (e.g. of three channel in total).

There are several embodiments for coupling the compensation unit to said drive field signal generator unit, said signal receiving unit and a drive-receiving coil. Generally, an inductive coupling or a capacitive coupling from the power amplifier of the drive field signal generator unit to the high current resonator (comprising the coupling unit, the measurement inductor and the drive-receiving coil) is used according to various embodiments. In particular, for inductive coupling a primary inductor unit is provided that is galvanically coupled to said drive field signal generator unit and inductively coupled to said measurement inductor. For capacitive coupling a coupling capacitor unit is provided that is coupled between said drive field signal generator unit, said measurement inductor and said drive-receiving coils. Inductive coupling has the advantage that it provides for a galvanic separation avoiding unwanted ground loops.

Further details of the interconnections between the elements of the compensation unit, the drive -receiving coils, the drive field signal generator unit and the signal receiving unit are defined in further dependent claims.

In another aspect of the present invention a coil arrangement for use in the proposed apparatus for influencing and/or detecting magnetic particles in a field of view is presented, which coil arrangement comprises:

one or more drive-receiving coils forming a bore for receiving a subject (e.g. a person or animal), a drive-receiving coil being configured both for changing the position in space of the two sub-zones in the field of view by means of a magnetic drive field so that the magnetization of the magnetic material changes locally and for acquiring detection signals, which detection signals depend on the magnetization in the field of view, which magnetization is influenced by the change in the position in space of the first and second sub- zone, and

a measurement inductor and a gradiometer inductor being inductively coupled, said measurement inductor and said gradiometer inductor being galvanically coupled to said one or more drive-receiving coils and being arranged outside said bore.

This other aspect of the invention can be combined with the optional or preferred embodiments of said first aspect of the invention.

According to still a further aspect of the present invention an apparatus for influencing and/or detecting magnetic particles in a field of view comprising such a coil arrangement is presented.

BRIEF DESCRIPTION OF THE DRAWINGS

These and other aspects of the invention will be apparent from and elucidated with reference to the embodiment(s) described hereinafter. In the following drawings

Fig. 1 shows a first embodiment of an MPI apparatus,

Fig. 2 shows an example of the selection field pattern produced by an apparatus as shown in Fig. 1,

Fig. 3 shows a second embodiment of an MPI apparatus,

Fig. 4 shows a third and a fourth embodiment of an MPI apparatus,

Fig. 5 shows a block diagram of an MPI apparatus according to the present invention,

Fig. 6 shows a block diagram of the general filtering scheme as used in an MPI apparatus,

Fig. 7 shows two circuit diagrams of an MPI apparatus according to the present invention comprising a capacitively coupled compensation unit,

Fig. 8 shows two circuit diagrams of an MPI apparatus according to the present invention comprising an inductively coupled compensation unit,

Fig. 9 shows a toroid comprising said measurement inductor and said gradiometer inductor, and

Fig. 10 shows two circuit diagrams of an auxiliary focus field generator unit.

DETAILED DESCRIPTION OF THE INVENTION

Before the details of the present invention shall be explained, basics of magnetic particle imaging shall be explained in detail with reference to Figs. 1 to 4. In particular, four embodiments of an MPI scanner for medical diagnostics will be described. An informal description of the data acquisition will also be given. The similarities and differences between the different embodiments will be pointed out. Generally, the present invention can be used in all these different embodiments of an MPI apparatus.

The first embodiment 10 of an MPI scanner shown in Fig. 1 has three pairs 12, 14, 16 of coaxial parallel circular coils, these coil pairs being arranged as illustrated in Fig. 1. These coil pairs 12, 14, 16 serve to generate the selection field as well as the drive and focus fields. The axes 18, 20, 22 of the three coil pairs 12, 14, 16 are mutually orthogonal and meet in a single point, designated the isocenter 24 of the MPI scanner 10. In addition, these axes 18, 20, 22 serve as the axes of a 3D Cartesian x-y-z coordinate system attached to the isocenter 24. The vertical axis 20 is nominated the y-axis, so that the x- and z-axes are horizontal. The coil pairs 12, 14, 16 are named after their axes. For example, the y-coil pair 14 is formed by the coils at the top and the bottom of the scanner. Moreover, the coil with the positive (negative) y-coordinate is called the y⁺-coil (y -coil), and similarly for the remaining coils. When more convenient, the coordinate axes and the coils shall be labelled with xi, x₂, and X3, rather than with x, y, and z.

The scanner 10 can be set to direct a predetermined, time-dependent electric current through each of these coils 12, 14, 16, and in either direction. If the current flows clockwise around a coil when seen along this coil's axis, it will be taken as positive, otherwise as negative. To generate the static selection field, a constant positive current I is made to flow through the z⁺-coil, and the current -I^s is made to flow through the z -coil. The z-coil pair 16 then acts as an anti-parallel circular coil pair.

It should be noted here that the arrangement of the axes and the nomenclature given to the axes in this embodiment is just an example and might also be different in other embodiments. For instance, in practical embodiments the vertical axis is often considered as the z-axis rather than the y-axis as in the present embodiment. This, however, does not generally change the function and operation of the device and the effect of the present invention.

The magnetic selection field, which is generally a magnetic gradient field, is represented in Fig. 2 by the field lines 50. It has a substantially constant gradient in the direction of the (e.g. horizontal) z-axis 22 of the z-coil pair 16 generating the selection field and reaches the value zero in the isocenter 24 on this axis 22. Starting from this field-free point (not individually shown in Fig. 2), the field strength of the magnetic selection field 50 increases in all three spatial directions as the distance increases from the field-free point. In a first sub-zone or region 52 which is denoted by a dashed line around the isocenter 24 the field strength is so small that the magnetization of particles present in that first sub-zone 52 is not saturated, whereas the magnetization of particles present in a second sub-zone 54 (outside the region 52) is in a state of saturation. In the second sub-zone 54 (i.e. in the residual part of the scanner's field of view 28 outside of the first sub-zone 52) the magnetic field strength of the selection field is sufficiently strong to keep the magnetic particles in a state of saturation.

By changing the position of the two sub-zones 52, 54 (including the field-free point) within the field of view 28 the (overall) magnetization in the field of view 28 changes. By determining the magnetization in the field of view 28 or physical parameters influenced by the magnetization, information about the spatial distribution of the magnetic particles in the field of view 28 can be obtained. In order to change the relative spatial position of the two sub-zones 52, 54 (including the field-free point) in the field of view 28, further magnetic fields, i.e. the magnetic drive field, and, if applicable, the magnetic focus field, are superposed to the selection field 50.

To generate the drive field, a time dependent current I°i is made to flow through both x-coils 12, a time dependent current I^D ₂ through both y-coils 14, and a time dependent current I^D ₃ through both z-coils 16. Thus, each of the three coil pairs acts as a parallel circular coil pair. Similarly, to generate the focus field, a time dependent current I i is made to flow through both x-coils 12, a current I ₂ through both y-coils 14, and a current p

I 3 through both z-coils 16.

It should be noted that the z-coil pair 16 is special: It generates not only its share of the drive and focus fields, but also the selection field (of course, in other embodiments, separate coils may be provided). The current flowing through the z^±-coil is I^D ₃

+ I F S The current flowing through the remaining two coil pairs 12, 14 is I D _¼ + I F

3 + I . k = 1, 2. Because of their geometry and symmetry, the three coil pairs 12, 14, 16 are well decoupled. This is wanted.

Being generated by an anti-parallel circular coil pair, the selection field is rotationally symmetric about the z-axis, and its z-component is nearly linear in z and independent of x and y in a sizeable volume around the isocenter 24. In particular, the selection field has a single field-free point (FFP) at the isocenter. In contrast, the contributions to the drive and focus fields, which are generated by parallel circular coil pairs, are spatially nearly homogeneous in a sizeable volume around the isocenter 24 and parallel to the axis of the respective coil pair. The drive and focus fields jointly generated by all three parallel circular coil pairs are spatially nearly homogeneous and can be given any direction and strength, up to some maximum strength. The drive and focus fields are also time- dependent. The difference between the focus field and the drive field is that the focus field varies slowly in time and may have a large amplitude, while the drive field varies rapidly and has a small amplitude. There are physical and biomedical reasons to treat these fields differently. A rapidly varying field with a large amplitude would be difficult to generate and potentially hazardous to a patient.

In a practical embodiment the FFP can be considered as a mathematical point, at which the magnetic field is assumed to be zero. The magnetic field strength increases with increasing distance from the FFP, wherein the increase rate might be different for different directions (depending e.g. on the particular layout of the device). As long as the magnetic field strength is below the field strength required for bringing magnetic particles into the state of saturation, the particle actively contributes to the signal generation of the signal measured by the device; otherwise, the particles are saturated and do not generate any signal.

The embodiment 10 of the MPI scanner has at least one further pair, preferably three further pairs, of parallel circular coils, again oriented along the x-, y-, and z- axes. These coil pairs, which are not shown in Fig. 1, serve as receive coils. As with the coil pairs 12, 14, 16 for the drive and focus fields, the magnetic field generated by a constant current flowing through one of these receive coil pairs is spatially nearly homogeneous within the field of view and parallel to the axis of the respective coil pair. The receive coils are supposed to be well decoupled. The time-dependent voltage induced in a receive coil is amplified and sampled by a receiver attached to this coil. More precisely, to cope with the enormous dynamic range of this signal, the receiver samples the difference between the received signal and a reference signal. The transfer function of the receiver is non-zero from zero Hertz ("DC") up to the frequency where the expected signal level drops below the noise level. Alternatively, the MPI scanner has no dedicated receive coils. Instead the drive field transmit coils are used as receive coils as is the case according to the present invention using combined drive-receiving coils.

The embodiment 10 of the MPI scanner shown in Fig. 1 has a cylindrical bore 26 along the z-axis 22, i.e. along the axis of the selection field. All coils are placed outside this bore 26. For the data acquisition, the patient (or object) to be imaged is placed in the bore 26 such that the patient's volume of interest - that volume of the patient (or object) that shall be imaged - is enclosed by the scanner's field of view 28 - that volume of the scanner whose contents the scanner can image. The patient (or object) is, for instance, placed on a patient table. The field of view 28 is a geometrically simple, isocentric volume in the interior of the bore 26, such as a cube, a ball, a cylinder or an arbitrary shape. A cubical field of view 28 is illustrated in Fig. 1.

The size of the first sub-zone 52 is dependent on the strength of the gradient of the magnetic selection field and on the field strength of the magnetic field required for saturation, which in turn depends on the magnetic particles. For a sufficient saturation of typical magnetic particles at a magnetic field strength of 80 A/m and a gradient (in a given space direction) of the field strength of the magnetic selection field amounting to 50x10 A/m , the first sub-zone 52 in which the magnetization of the particles is not saturated has dimensions of about 1 mm (in the given space direction).

The patient's volume of interest is supposed to contain magnetic nanoparticles. Prior to the diagnostic imaging of, for example, a tumor, the magnetic particles are brought to the volume of interest, e.g. by means of a liquid comprising the magnetic particles which is injected into the body of the patient (object) or otherwise administered, e.g. orally, to the patient.

Generally, various ways for bringing the magnetic particles into the field of view exist. In particular, in case of a patient into whose body the magnetic particles are to be introduced, the magnetic particles can be administered by use of surgical and non-surgical methods, and there are both methods which require an expert (like a medical practitioner) and methods which do not require an expert, e.g. can be carried out by laypersons or persons of ordinary skill or the patient himself / herself. Among the surgical methods there are potentially non-risky and/or safe routine interventions, e.g. involving an invasive step like an injection of a tracer into a blood vessel (if such an injection is at all to be considered as a surgical method), i.e. interventions which do not require considerable professional medical expertise to be carried out and which do not involve serious health risks. Further, nonsurgical methods like swallowing or inhalation can be applied.

Generally, the magnetic particles are pre-delivered or pre-administered before the actual steps of data acquisition are carried out. In embodiments, it is, however, also possible that further magnetic particles are delivered / administered into the field of view.

An embodiment of magnetic particles comprises, for example, a spherical substrate, for example, of glass which is provided with a soft-magnetic layer which has a thickness of, for example, 5 nm and consists, for example, of an iron-nickel alloy (for example, Permalloy). This layer may be covered, for example, by means of a coating layer which protects the particle against chemically and/or physically aggressive environments, e.g. acids. The magnetic field strength of the magnetic selection field 50 required for the saturation of the magnetization of such particles is dependent on various parameters, e.g. the diameter of the particles, the used magnetic material for the magnetic layer and other parameters.

In the case of e.g. a diameter of 10 μιη with such magnetic particles, a magnetic field of approximately 800 A/m (corresponding approximately to a flux density of 1 mT) is then required, whereas in the case of a diameter of 100 μιη a magnetic field of 80 A/m suffices. Even smaller values are obtained when a coating of a material having a lower saturation magnetization is chosen or when the thickness of the layer is reduced.

In practice, magnetic particles commercially available under the trade name Resovist (or similar magnetic particles) are often used, which have a core of magnetic material or are formed as a massive sphere and which have a diameter in the range of nanometers, e.g. 40 or 60 nm.

For further details of the generally usable magnetic particles and particle compositions, the corresponding parts of EP 1224542, WO 2004/091386, WO 2004/091390, WO 2004/091394, WO 2004/091395, WO 2004/091396, WO 2004/091397, WO 2004/091398, WO 2004/091408 are herewith referred to, which are herein incorporated by reference. In these documents more details of the MPI method in general can be found as well.

During the data acquisition, the x-, y-, and z-coil pairs 12, 14, 16 generate a position- and time-dependent magnetic field, the applied field. This is achieved by directing suitable currents through the field generating coils. In effect, the drive and focus fields push the selection field around such that the FFP moves along a preselected FFP trajectory that traces out the volume of scanning - a superset of the field of view. The applied field orientates the magnetic nanoparticles in the patient. As the applied field changes, the resulting magnetization changes too, though it responds nonlinearly to the applied field. The sum of the changing applied field and the changing magnetization induces a time-dependent voltage V_k across the terminals of the receive coil pair along the Xk-axis. The associated receiver converts this voltage to a signal S_k, which it processes further. Like the first embodiment 10 shown in Fig. 1, the second embodiment 30 of the MPI scanner shown in Fig. 3 has three circular and mutually orthogonal coil pairs 32, 34, 36, but these coil pairs 32, 34, 36 generate the selection field and the focus field only. The z- coils 36, which again generate the selection field, are filled with ferromagnetic material 37. The z-axis 42 of this embodiment 30 is oriented vertically, while the x- and y-axes 38, 40 are oriented horizontally. The bore 46 of the scanner is parallel to the x-axis 38 and, thus, perpendicular to the axis 42 of the selection field. The drive field is generated by a solenoid (not shown) along the x-axis 38 and by pairs of saddle coils (not shown) along the two remaining axes 40, 42. These coils are wound around a tube which forms the bore. The drive field coils also serve as receive coils.

To give a few typical parameters of such an embodiment: The z-gradient of the selection field, G, has a strength of G/μο = 2.5 T/m, where μο is the vacuum permeability. The temporal frequency spectrum of the drive field is concentrated in a narrow band around 25 kHz (up to approximately 150 kHz). The useful frequency spectrum of the received signals lies between 50 kHz and 1 MHz (eventually up to approximately 15 MHz). The bore has a diameter of 120 mm. The biggest cube 28 that fits into the bore 46 has an edge length

Since the construction of field generating coils is generally known in the art, e.g. from the field of magnetic resonance imaging, this subject need not be further elaborated herein.

In an alternative embodiment for the generation of the selection field, permanent magnets (not shown) can be used. In the space between two poles of such (opposing) permanent magnets (not shown) there is formed a magnetic field which is similar to that shown in Fig. 2, that is, when the opposing poles have the same polarity. In another alternative embodiment, the selection field can be generated by a mixture of at least one permanent magnet and at least one coil.

Fig. 4 shows two embodiments of the general outer layout of an MPI apparatus 200, 300. Fig. 4A shows an embodiment of the proposed MPI apparatus 200 comprising two selection-and-focus field coil units 210, 220 which are basically identical and arranged on opposite sides of the examination area 230 formed between them. Further, a drive field coil unit 240 is arranged between the selection-and-focus field coil units 210, 220, which are placed around the area of interest of the patient (not shown). The selection-and- focus field coil units 210, 220 comprise several selection-and-focus field coils for generating a combined magnetic field representing the above-explained magnetic selection field and magnetic focus field. In particular, each selection-and-focus field coil unit 210, 220 comprises a, preferably identical, set of selection-and-focus field coils. Details of said selection-and-focus field coils will be explained below.

The drive field coil unit 240 comprises a number of drive field coils for generating a magnetic drive field. These drive field coils may comprise several pairs of drive field coils, in particular one pair of drive field coils for generating a magnetic field in each of the three directions in space. In an embodiment the drive field coil unit 240 comprises two pairs of saddle coils for two different directions in space and one solenoid coil for generating a magnetic field in the longitudinal axis of the patient.

The selection-and-focus field coil units 210, 220 are generally mounted to a holding unit (not shown) or the wall of room. Preferably, in case the selection-and-focus field coil units 210, 220 comprise pole shoes for carrying the respective coils, the holding unit does not only mechanically hold the selection-and-focus field coil unit 210, 220 but also provides a path for the magnetic flux that connects the pole shoes of the two selection-and- focus field coil units 210, 220.

As shown in Fig. 4a, the two selection-and-focus field coil units 210, 220 each include a shielding layer 211, 221 for shielding the selection-and-focus field coils from magnetic fields generated by the drive field coils of the drive field coil unit 240.

In the embodiment of the MPI apparatus 201 shown in Fig. 4B only a single selection-and-focus field coil unit 220 is provided as well as the drive field coil unit 240. Generally, a single selection-and-focus field coil unit is sufficient for generating the required combined magnetic selection and focus field. Said single selection-and-focus field coil unit 220 may thus be integrated into a (not shown) patient table on which a patient is placed for the examination. Preferably, the drive field coils of the drive field coil unit 240 may be arranged around the patient's body already in advance, e.g. as flexible coil elements. In another implementation, the drive field coil unit 240 can be opened, e.g. separable into two subunits 241, 242 as indicated by the separation lines 243, 244 shown in Fig. 4b in axial direction, so that the patient can be placed in between and the drive field coil subunits 241, 242 can then be coupled together.

In still further embodiments of the MPI apparatus, even more selection-and- focus field coil units may be provided which are preferably arranged according to a uniform distribution around the examination area 230. However, the more selection-and-focus field coil units are used, the more will the accessibility of the examination area for placing a patient therein and for accessing the patient itself during an examination by medical assistance or doctors be limited.

Fig. 5 shows a general block diagram of an MPI apparatus 100 according to the present invention. The general principles of magnetic particle imaging explained above are valid and applicable to this embodiment as well, unless otherwise specified.

The embodiment of the apparatus 100 shown in Fig. 5 comprises various coils for generating the desired magnetic fields. First, the coils and their functions in MPI shall be explained.

For generating the combined magnetic selection-and-focus field, selection- and-focus device 110 are provided. The magnetic selection-and-focus field has a pattern in space of its magnetic field strength such that the first sub-zone (52 in Fig. 2) having a low magnetic field strength where the magnetization of the magnetic particles is not saturated and a second sub-zone (54 in Fig. 4) having a higher magnetic field strength where the magnetization of the magnetic particles is saturated are formed in the field of view 28, which is a small part of the examination area 230, which is conventionally achieved by use of the magnetic selection field. Further, by use of the magnetic selection-and-focus field the position in space of the field of view 28 within the examination area 230 can be changed, as conventionally done by use of the magnetic focus field.

The selection-and-focus device 110 comprises at least one set of selection- and-focus field coils 114 and a selection-and-focus field generator unit 112 for generating selection-and-focus field currents to be provided to said at least one set of selection-and- focus field coils 114 (representing one of the selection-and-focus field coil units 210, 220 shown in Figs. 4A, 4B) for controlling the generation of said magnetic selection-and-focus field. Preferably, a separate generator subunit is provided for each coil element (or each pair of coil elements) of the at least one set of selection-and-focus field coils 114. Said selection- and-focus field generator unit 112 comprises a controllable current source (generally including an amplifier) and a filter unit which provide the respective coil element with the field current to individually set the gradient strength and field strength of the contribution of each coil to the magnetic selection-and-focus field. It shall be noted that the filter unit 114 can also be omitted. Further, separate focus and selection device are provided in other embodiments.

For generating the magnetic drive field the apparatus 100 further comprises drive device 120 comprising a drive field signal generator unit 122 and a set of drive field coils 124 (representing the drive coil unit 240 shown in Figs. 4A, 4B) for changing the position in space and/or size of the two sub-zones in the field of view by means of a magnetic drive field so that the magnetization of the magnetic material changes locally. As mentioned above said drive field coils 124 preferably comprise two pairs 125, 126 of oppositely arranged saddle coils and one solenoid coil 127. Other implementations, e.g. three pairs of coil elements, are also possible.

The drive field signal generator unit 122 preferably comprises a separate drive field signal generation subunit for each coil element (or at least each pair of coil elements) of said set of drive field coils 124. Said drive field signal generator unit 122 preferably comprises a drive field current source (preferably including a power amplifier) and a filter unit for providing a time-dependent drive field current to the respective drive field coil.

The selection-and-focus field signal generator unit 112 and the drive field signal generator unit 122 are preferably controlled by a control unit 150, which preferably controls the selection-and-focus field signal generator unit 112 such that the sum of the field strengths and the sum of the gradient strengths of all spatial points of the selection field is set at a predefined level. For this purpose the control unit 150 can also be provided with control instructions by a user according to the desired application of the MPI apparatus, which, however, is preferably omitted according to the present invention.

For using the MPI apparatus 100 for determining the spatial distribution of the magnetic particles in the examination area (or a region of interest in the examination area), particularly to obtain images of said region of interest, signal detection receiving device, in particular a receiving coil, and a signal receiving unit 140, which receives signals detected by said receiving device, are provided. Conventionally, one to three separate receiving coils are provided in an MPI apparatus as receiving device. According to the present invention, however, one to three of said drive field coils 124 (or drive field coil pairs) act (simultaneously or alternately) as receiving coils for receiving detection signals. Accordingly, these drive field coils are called "drive-receiving coils" herein.

The generation of magnetic drive fields and the detection of detection signals can be performed simultaneously or alternately. Preferably, all three drive-receiving coils (or coil pairs) 125, 126, 127 act as receiving coils and three receiving units 140 - one per drive- receiving coil (or coil pair) - are provided in practice, but more than three drive-receiving coils and receiving units can be also used, in which case the acquired detection signals are not 3-dimensional but K-dimensional, with K being the number of drive-receiving coils.

Said signal receiving unit 140 comprises a filter unit 142 for filtering the received detection signals. The aim of this filtering is to separate measured values, which are caused by the magnetization in the examination area which is influenced by the change in position of the two part-regions (52, 54), from other, interfering signals (in particular crosstalk of the fundamental frequency). To this end, the filter unit 142 may be designed for example such that signals which have temporal frequencies that are smaller than the temporal frequencies with which the drive-receiving coil(s) is (are) operated, or smaller than twice these temporal frequencies, do not pass the filter unit 142. The signals are then transmitted via an amplifier unit 144 to an analog/digital converter 146 (ADC).

The digitized signals produced by the analog/digital converter 146 are fed to an image processing unit (also called reconstruction device) 152, which reconstructs the spatial distribution of the magnetic particles from these signals and the respective position which the first part-region 52 of the first magnetic field in the examination area assumed during receipt of the respective signal and which the image processing unit 152 obtains from the control unit 150. The reconstructed spatial distribution of the magnetic particles is finally transmitted via the control device 150 to a computer 154, which displays it on a monitor 156. Thus, an image can be displayed showing the distribution of magnetic particles in the field of view of the examination area.

In other applications of the MPI apparatus 100, e.g. for influencing the magnetic particles (for instance for a hyperthermia treatment) or for moving the magnetic particles (e.g. attached to a catheter for moving the catheter or attached to a medicament for moving the medicament to a certain location) the receiving device may also be omitted or simply not used.

Further, an input unit 158 may optionally be provided, for example a keyboard. A user may therefore be able to set the desired direction of the highest resolution and in turn receives the respective image of the region of action on the monitor 156. If the critical direction, in which the highest resolution is needed, deviates from the direction set first by the user, the user can still vary the direction manually in order to produce a further image with an improved imaging resolution. This resolution improvement process can also be operated automatically by the control unit 150 and the computer 154. The control unit 150 in this embodiment sets the gradient field in a first direction which is automatically estimated or set as start value by the user. The direction of the gradient field is then varied stepwise until the resolution of the thereby received images, which are compared by the computer 154, is maximal, respectively not improved anymore. The most critical direction can therefore be found respectively adapted automatically in order to receive the highest possible resolution.

Still further, according to the present invention a compensation unit 160 is provided comprising a measurement inductor and a gradiometer inductor. Said measurement inductor and said gradiometer inductor are inductively coupled, and said compensation unit 160 is electrically coupled between said drive field signal generator unit 122, said signal receiving unit 140 and said drive-receiving coils 124, in particular the drive-receiving coil to which a particular compensation unit is associated (generally, for each drive-receiving coil a separate compensation unit may be provided). This will be explained in more detail below.

Fig. 6 shows a block diagram of the general filtering scheme 300 as used in an MPI apparatus and spectra of various signals taken at different connections in said filtering scheme. As explained above MPI is based on the detection of harmonics as generated by magnetic particles subjected to an external sinusoidal magnetic field excitation by use of a synthesizer 301 and a power amplifier 302, which (together with the band pass filter 303) basically represent the drive field signal generator unit 122 shown in Fig. 5. Excitation and reception (by use of a low-noise amplifier 308 and an ADC 309, which (together with the band stop filter 307) basically represent the signal receiving unit 140 shown in Fig. 5, are taking place simultaneously, and are solely separated in the frequency domain. The classic separation is realized by notch filters (e.g. LC resonators), i.e. a band pass filter 303 in front of the transmit coil 304 (drive field coil) and a band stop filter 307 after the receive coil 306, wherein said transmit coil 304 and said receive coil 306 are separate coils arranged close to the bore 305 in which the patient is placed for examination. Due to the higher sensitivity of coils that are nearest to the patient, there is a competition between the transmit coil 304 and the receive coil 306 on the space very near around the patient. This "competition" is solved according to the present invention by using a joint transmit / receive coil, i.e. a drive- receiving coil as mentioned above.

Since there is crosstalk from the transmit side to the receive side, in particular harmonics generated from the power amplifier 302, a lot of effort is spent on the band pass filter 303 to ensure that no harmonics from the drive field enter the receive path. However, it was found, that the success of this effort is limited finally by the component of the filter 303 itself: particularly the capacitors (but also other components and materials) behave non- linearily. The degree to which they are non-linear is so small that it is hard to measure and it seems to be of no concern to other applications. Nevertheless it becomes limiting for this reception scheme, despite efforts to identify an optimum capacitor technology.

Fig. 7 shows two circuit diagrams of an MPI apparatus 400, 400' according to the present invention comprising a capacitively coupled compensation unit 320. Same elements as in the MPI apparatus 300 shown in Fig. 6 are provided with like reference numbers.

The MPI apparatus 400 shown in Fig. 7A has an electrically asymmetric layout. It comprises a joint transmit and receive coil 330 (also indicated by L_D and referred to as drive-receiving coil), which corresponds to one of the coils (or coil pairs 125, 126, 127 shown in Fig. 5). There is electrically only one drive-receiving coil 330 used for both transmission of fundamental frequency and for reception of harmonic frequencies. Hence, there are no differences in sensitivity or drift between different components.

The compensation unit 320 comprises a measurement inductor 321 (also indicated by L_M) and a gradiometer inductor 322 (also indicated by L_G), which are inductively coupled like in a transformer. Said compensation unit 320 is electrically coupled between said drive field signal generator unit (122 in Fig. 5; represented by elements 302, 303 in Fig. 7), said signal receiving unit (140 in Fig. 5; represented by elements 307, 308 in Fig. 7) and the drive-receiving coil 330.

In this embodiment first end terminals 321a, 322a of said measurement inductor 321 and said gradiometer inductor 322 are coupled to a first end terminal 330a of the associated drive-receiving coil 330. Further, a second end terminal 330b of said drive- receiving coil 330 is coupled to a second end terminal 122b of said drive field signal generator unit 122 and a second end terminal 140b of said signal receiving unit 140, which are commonly coupled to ground. Still further, a second end terminal 321b of said measurement inductor 321 is coupled to a first end terminal 122a of said drive field signal generator unit 122 and a second end terminal 322b of said gradiometer inductor 322 is coupled to a first end terminal 140a of said signal receiving unit 140.

The drive field signal generator unit 122 further comprises a primary capacitor unit 340 for capacitively coupling said drive field signal generator unit 122 to said inductive coupling unit 320 and said drive-receiving coil 330, which also provides for an impedance transformation. Said primary capacitor unit 340 comprises two primary capacitors 341, 342, wherein the first primary capacitor 341 is coupled to the output terminals of said band pass filter 303 and the second primary capacitor 342 is coupled between said first primary capacitor 341 and the end terminal 321b of said measurement inductor 321 . The primary capacitor unit 340, the measurement inductor 321 and the drive-receiving coil 330 form a high current resonator into which the energy is coupled by said primary capacitor unit 340.

The MPI apparatus 400' shown in Fig. 7B has an electrically symmetric layout. Besides the (first) compensation unit 320 comprising a first measurement inductor 321 (also indicated by L_MI) and a first gradiometer inductor 322 (also indicated by L_GI) a (second) compensation unit 320' comprising a second measurement inductor 321 ' (also indicated by L_M2) and a second gradiometer inductor 322' (also indicated by LQ₂) is provided between the drive-receiving coil 330, the second end terminal 122b of said drive field signal generator unit 122 and the second end terminal 140b of said signal receiving unit 140.

The primary capacitor unit 340' comprises four symmetrically arranged primary capacitors. The first primary capacitor 341 of the primary capacitor unit 340 is split into two primary capacitors 341a, 341b, between which a ground potential is coupled. Further, the primary capacitors 342, 343 are coupled between a respective output of the filter 303 and the respective end terminal 122a, 122b.

The symmetric arrangement generally has, compared to the asymmetric arrangement, a better common-mode suppression and is hence less sensitive to external interferences.

A gradiometric cancellation scheme is thus used according to the present invention to suppress harmonic background in the receive path. The measurement inductor 320 and the gradiometer inductor 321 form a compensation unit and are coupled like a transformer. In the drive-receiving coil 330 not only the (desired) detection signal is coupled but also the excitation signal (drive signal). The excitation signal is also coupled into the measurement inductor 320 (which receives the same current from the drive field signal generator unit 122 as the drive-receiving coil 330. By the coupling (preferably by substantially 1:1) of the measurement inductor 320 to the gradiometer inductor 321 substantially the same voltage is induced into the gradiometer inductor 321, but with different polarity. This has the effect that in the receive path (formed by the drive-receiving coil 330 and the gradiometer inductor 322) the excitation signal coupled into the drive-receiving coil 330 is counter-balanced by the excitation signal coupled by the measurement inductor 320 into the gradiometer inductor 321 with reversed polarity. Thus, at the signal receiving unit 140 only the desired detection signal measured by the drive-receiving coils and caused by the magnetic particles is registered. By use of a combined coil as drive-receiving coil 330 that fulfils both functions of the (conventionally separate) drive coil and receiving coil space is saved in the bore of the imaging volume into which the patient is placed for an examination. Generally only a single drive-receiving coil 330 is used, but in other embodiments one drive-receiving coil per direction is used, or one drive-receiving coil for one direction and two (conventional) drive coils for the other directions are used. Thus, compared to known solutions less hardware elements are required and more space within the bore is available. The measurement inductor 320 and the gradiometer inductor 321 can be placed outside the bore and thus do not consume any space into which the patient is to be placed. Fig. 8 shows two circuit diagrams of an MPI apparatus 500, 500' according to the present invention comprising an inductively coupled compensation unit 320. Same elements as in the MPI apparatus 300 shown in Fig. 6 are provided with like reference numbers.

The MPI apparatus 500 shown in Fig. 8A has an electrically asymmetric layout. In this embodiment the first end terminal 321a of said measurement inductor 321 is, via a secondary capacitor 360, coupled to a first end terminal 330a of the associated drive- receiving coil 330. A second end terminal 321b of said measurement inductor 321 are coupled to a first second terminal 140b of said signal receiving unit 140 and a second end terminal 330b of said drive-receiving coil 330. A first end terminal 322b of said gradient inductor 322 is coupled to a first end terminal 140a of said signal receiving unit 140. A second end terminal 322a of said gradient inductor 322 is coupled to a first end terminal 330a of the drive-receiving coil 330.

Alternatively, the secondary capacitor 360 can be located at the other side of the measurement inductor 321 (i.e. Indeed, as both components are electrically in series, it makes no difference), the measurement inductor 321 being then located between the capacitor 360 and the drive field coil 330 (not depicted in any figure) - the connection of the second end terminal 322a of the gradient gradiometer 322 being still coupled to a first end terminal 330a of the drive-receiving coil 330, between the assembly capacitor 360/measurement inductor 321 and the drive field coil 330. Like in the MPI apparatus 400, 400' the combination of measurement inductor 321 and gradient inductor 322 can be regarded as compensation unit 320.

The drive field signal generator unit 122 further comprises a primary inductor unit 350, in particular a single inductor (also referred to as Lp), that is galvanically coupled to the end terminals 122a, 122b of said drive field signal generator unit 122 and inductively coupled to said measurement inductor 321. The primary inductor unit 350 together with said measurement inductor 321 also provides for an impedance transformation. The measurement inductor 321, the secondary capacitor 360 and the drive-receiving coil 330 form a high- current resonator into which the energy is coupled by said primary inductor unit 350.

The MPI apparatus 500' shown in Fig. 8B has an electrically symmetric layout. Besides the (first) compensation unit 320 comprising a first measurement inductor 321 (also indicated by L_Mi) and a first gradiometer inductor 322 (also indicated by LQI) a (second) compensation unit 320' comprising a second measurement inductor 321 ' (also indicated by L_M2) and a second gradiometer inductor 322' (also indicated by L_G2) is provided. Between the first and second measurement inductors 321, 321 ' a ground potential is coupled. The other end terminals are coupled, via a respective secondary capacitor 360, 360' to the drive-receiving coil 330. The gradiometer inductors 322, 322' are coupled between a respective end terminal 140a, 140b of the signal receiving unit 140 and a respective end terminal 330a, 330b of the drive-receiving coil 330. Alternatively, either or the two secondary capacitors 360, 360' are located, respectively, at the other side of the measurement inductors 321, 321 ' (i.e. indeed, as both components are electrically in series, it makes no difference), the measurement inductors 321, 321 ' being then respectively located between the capacitor 360, 360' and the drive field coil 330 (not depicted in any figure).

Fig. 9 shows a preferred embodiment of the compensation unit 320 according to which the measurement inductor 32 land the gradiometer inductor 322 are wound like a toroid. The windings of the measurement inductor 321 and the gradiometer inductor 322 are preferably wound alternatingly. The core 323 of the toroid is preferably air.

Preferably, according to the present invention a cancellation condition should be fulfilled as follows:

wherein the coupling k between the measurement inductor L_M and the gradiometer inductor L_G is defined b

and M represents the mutual inductance.

This is derived from the following assumptions:

U _RX = !_TX (j^Lo - j»M) + U_NP

wherein U_Rx is the voltage signal across the signal receiving unit 140, Ιχ_χ is the current in the high-current resonator, and U _P is the signal induced by the magnetic particles into the drive- receiving coil 330.

In initial implementations of pre-clinical, i.e. small-animal MPI scanners, the amplitude and frequency of the drive field thus far have been around 20mT peak and 25 kHz, respectively. However, it has been found that the signal amplitude of the drive field should preferably not be larger than 5-10 mT to avoid any negative effects on the patient tissue, such as a stimulation of muscle tissue or even a heating of tissue. Hence, to substantially maintain the voxel acquisition rate the drive field frequency for clinical, i.e. human-size MPI scanners, should be increased from approx. 25 kHz to, for instance, approx. 150 kHz. The result is that the size of the field of view 28 within the examination area 230 is reduced which, in turn, means that scanning the region of interest requires more time if the (smaller) field of view 28 is moved through the region of interest by use of the same focus field (or selection-and-focus field) having a rather low frequency, e.g. in the range of 10 Hz.

Hence, it is proposed in an embodiment to make use of an auxiliary magnetic focus field having a larger frequency (e.g. in the range from 25 to 200 Hz, preferably around 100 Hz) than the already available (standard) magnetic focus field in order to move the (now smaller) field of view 28 faster and, thus, to scan the region of interest in substantially the same time.

In an embodiment auxiliary focus field coil(s) (preferably one focus field coil or coil pair per direction) are provided in addition to the coils of the MPI apparatus, as e.g. schematically shown in Fig. 5. In another embodiment the drive-receiving coils (or, in other MPI apparatus, the drive field coils) are used for generating this (these) auxiliary magnetic focus field(s).

For the latter embodiment that is preferably used in the MPI apparatus according to the present invention an auxiliary focus field generator unit is provided as shown in Fig. 10. Fig. 10A shows a circuit diagram of an auxiliary focus field generator unit 170 for use with the asymmetric arrangements of the MPI apparatus 400, 500 shown in Figs. 7A, 8A. It comprises a power amplifier 171, an (asymmetric) filter 172 and two terminals 173, 174 through which the auxiliary focus field generator unit 170 is preferably coupled to the end terminals 140a, 140b of the signal receiving unit 140. This has the advantage that the available gradient inductor 322 (L_G) favorably avoids (negative and unwanted) back couplings from the magnetic drive field into the power amplifier 171 of the auxiliary focus field generator unit 170. Fig. 10B shows a circuit diagram of an auxiliary focus field generator unit 170' for use with the symmetric arrangements of the MPI apparatus 400', 500' shown in Figs. 7B, 8B. Like the auxiliary focus field generator unit 170 it comprises a power amplifier 171, an optional (asymmetric) filter 172 and two terminals 173, 174 through which the auxiliary focus field generator unit 170' is preferably coupled to the end terminals 140a, 140b of the signal receiving unit 140. Further, between the filter 172 and the two terminals 173, 174 an insulating transformer 175 and a symmetric filter 176 are optionally provided for symmetrization.

In summary, according to the present invention the space and the copper cross- section of the drive-receiving coil(s) is optimised. If electrically separate coils were used, and the space were given and considered as 100 %, then one coil could use p %, and the other coil could only use (1-p) %. Therefore, both coils would have less copper available, and both coils would have higher resistance. Resistance, however, is undesired since it leads to higher losses (and more cooling efforts) for the transmit coil, and to more noise (and hence reduced SNR) for the receive coil.

Further, according to the present invention no reduction of bore size is necessary. The full size of the bore is available for the patient. This is the same argument as above, but assuming both coils would like to have 100% space, then this sums up to 200%, so this space must be sacrificed elsewhere. Either the patient bore is reduced, or the gap between the selection/focus field coils is increased, both of which is not beneficial.

Still further, compensation is provided in combination with an external measurement inductor. The gradiometer is hence dislocated from the bore, the field of view and the magnetic particles, whereas its effect is however kept and realised by a transformer. This transformer can be at any position, where space is available, so it can be realised as a larger unit with more copper, leading to reduced resistance (and hence losses/noise).

There are various trade-offs with respect to the ratio of drive field and receiver coil to measurement coil, but generally they can be considered to have about identical inductances. Therefore, in order to fulfil the cancellation condition, and not to make gradiometer coil too large (which again is detrimental), the magnitude of the coupling coefficient k needs to be large, i.e. strong coupling is preferred. The parasitic capacitance of the transformer (not drawn) shall nevertheless be minimal.

General features of the proposed gradiometer are a wideband decoupling that stops fundamental and harmonics from all (so far typically three) drive fields, reduced requirements on the band pass filter in the transmit path (simpler topology, less weight, volume, losses, and cost), and reduced requirements on linearity of capacitors within the band pass filter and the high-current resonator towards the transmit coil.

While the invention has been illustrated and described in detail in the drawings and foregoing description, such illustration and description are to be considered illustrative or exemplary and not restrictive; the invention is not limited to the disclosed embodiments. Other variations to the disclosed embodiments can be understood and effected by those skilled in the art in practicing the claimed invention, from a study of the drawings, the disclosure, and the appended claims.

In the claims, the word "comprising" does not exclude other elements or steps, and the indefinite article "a" or "an" does not exclude a plurality. A single element or other unit may fulfill the functions of several items recited in the claims. The mere fact that certain measures are recited in mutually different dependent claims does not indicate that a combination of these measures cannot be used to advantage.

Any reference signs in the claims should not be construed as limiting the scope.

Claims

CLAIMS:

1. An apparatus (100) for influencing and/or detecting magnetic particles in a field of view (28), which apparatus comprises:

selection device comprising a selection field signal generator unit (110) and selection field elements (116) for generating a magnetic selection field (50) having a pattern in space of its magnetic field strength such that a first sub-zone (52) having a low magnetic field strength where the magnetization of the magnetic particles is not saturated and a second sub-zone having a higher magnetic field strength where the magnetization of the magnetic particles is saturated are formed in the field of view (28),

drive and receiving device comprising a drive field signal generator unit (122), a signal receiving unit (140) and one or more drive-receiving coils (330), a drive-receiving coil being configured both for changing the position in space of the two sub-zones (52, 54) in the field of view (28) by means of a magnetic drive field so that the magnetization of the magnetic material changes locally and for acquiring detection signals, which detection signals depend on the magnetization in the field of view (28), which magnetization is influenced by the change in the position in space of the first and second sub-zone (52, 54), and

a compensation unit (320) comprising a measurement inductor (321) and a gradiometer inductor (322), said measurement inductor (321) and said gradiometer inductor (322) being inductively coupled, wherein said compensation unit (320) is electrically coupled to said drive field signal generator unit (122), said signal receiving unit (140) and a drive- receiving coil (330).

2. The apparatus (100) as claimed in claim 1,

wherein first end terminals (321a, 322a) of said measurement inductor (321) and said gradiometer inductor (322) are coupled to a first or second end terminal (330a; 330b) of an associated drive-receiving coil (330).

3. The apparatus (100) as claimed in claim 2,

wherein a second end terminal (330b) of said drive-receiving coil (330) is coupled to a second end terminal (122b) of said drive field signal generator unit (122) and a second end terminal (140b) of said signal receiving unit (140).

4. The apparatus (100) as claimed in claim 2 or 3,

wherein a second end terminal (321b) of said measurement inductor (321) is coupled to a first end terminal (122a) of said drive field signal generator unit (122) and a second end terminal (322b) of said gradiometer inductor (322) is coupled to a first end terminal (140a) of said signal receiving unit (140).

5. The apparatus (100) as claimed in claim 2,

wherein a first end terminal (330a) of said drive-receiving coil (330) is coupled, via a secondary capacitor (360) to a first end terminal (321a) of said measurement inductor (321), a secondary capacitor (360) is coupled to the measurement inductor (321), and the first end terminal (140a) of said signal receiving unit (140) is coupled, via said gradiometer inductor (322) to a first end terminal (330a) of said drive-receiving coil (330).

6. The apparatus (100) as claimed in claim 2 or 5,

wherein a second end terminal (321b) of said measurement inductor (321) is coupled to a second end terminal (140b) of said signal receiving unit (140), a second end terminal (330b) of said drive-receiving coil (330) and optionally a second end terminal (122b) of said drive field signal generator unit (122).

7. The apparatus (100) as claimed in claim 2,

comprising a first compensation unit (320) and a second compensation unit (320') symmetrically coupled to said drive-receiving coil (330).

8. The apparatus (100) as claimed in claim 1,

further comprising a primary inductor unit (350) galvanically coupled to said drive field signal generator unit (122) and inductively coupled to said measurement inductor (321).

9. The apparatus (100) as claimed in claim 1,

further comprising a primary capacitor unit (340) coupled between said drive field signal generator unit (122), said measurement inductor (321) and said drive-receiving coils (330).

10. The apparatus (100) as claimed in claim 1,

wherein said compensation unit (320) is physically arranged separate from said drive- receiving coils (330), at a distant location from said drive-receiving coils (330) and/or outside a bore formed by said drive-receiving coils (330).

11. The apparatus (100) as claimed in claim 1,

wherein said drive-receiving coils (330) are additionally configured to generate an auxiliary magnetic focus field having a frequency in the range from 25 to 200 Hz, in particular from 40 to 120 Hz, for moving the field of view (28),

wherein said apparatus further comprises an auxiliary focus field generator unit (170, 170').

12. The apparatus (100) as claimed in claim 1,

wherein said measurement inductor (321) and said gradiometer inductor (322) are wound as toroid.

13. The apparatus (100) as claimed in claim 1,

wherein said compensation unit (320) comprises a measurement inductor (321) and a gradiometer inductor (322) per drive-receiving coil (330), said measurement inductor (321) and said gradiometer inductor (322) being inductively coupled.

14. A method for influencing and/or detecting magnetic particles in a field of view (28), which method comprises the steps of:

generating a magnetic selection field (50) having a pattern in space of its magnetic field strength such that a first sub-zone (52) having a low magnetic field strength where the magnetization of the magnetic particles is not saturated and a second sub -zone having a higher magnetic field strength where the magnetization of the magnetic particles is saturated are formed in the field of view (28),

changing the position in space of the two sub-zones (52, 54) in the field of view (28) by means of a magnetic drive field so that the magnetization of the magnetic material changes locally by means of one or more drive-receiving coils (330),

acquiring detection signals by means of at least one drive-receiving coil (330), which detection signals depend on the magnetization in the field of view (28), which magnetization is influenced by the change in the position in space of the first and second sub- zone (52, 54), and

coupling detection signals onto a compensation unit (320) comprising a measurement inductor (321) and a gradiometer inductor (322), said measurement inductor (321) and said gradiometer inductor (322) being inductively coupled, wherein said compensation unit (320) is electrically coupled to said drive field signal generator unit (122), said signal receiving unit (140) and a drive-receiving coil (330).

15. Computer program comprising program code elements for causing a computer to control an apparatus as claimed in claim 1 to carry out the steps of the method as claimed in claim 14 when said computer program is carried out on the computer.

16. A coil arrangement for use in the proposed apparatus for influencing and/or detecting magnetic particles in a field of view is presented, which coil arrangement comprises:

one or more drive-receiving coils (330) forming a bore for receiving a subject, a drive-receiving coil being configured both for changing the position in space of the two sub-zones in the field of view by means of a magnetic drive field so that the magnetization of the magnetic material changes locally and for acquiring detection signals, which detection signals depend on the magnetization in the field of view, which magnetization is influenced by the change in the position in space of the first and second sub-zone, and

a measurement inductor (320) and a gradiometer inductor (321) being inductively coupled, said measurement inductor and said gradiometer inductor being galvanically coupled to said one or more drive-receiving coils and being arranged outside said bore.