Patent Publication Number: US-2017354344-A1

Title: Cabling arrangment, coil apparatus and apparatus for influencing and/or detecting magnetic particles

Description:
FIELD OF THE INVENTION 
     The present invention relates to a cabling arrangement and a coil apparatus, in particular for use in a magnetic particle imaging apparatus, as well as an apparatus and method for influencing and/or detecting magnetic particles in a field of view. 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&#39;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 (in the following reference is mostly made to the field-free point, which shall however include the option of using a field-free line instead). 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&#39;s volume of interest is enclosed by the scanner&#39;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 A1 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. 
     An MPI apparatus and method 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 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. 
     Within MPI, image information is gathered by analysing the weak non-linearity of the magnetic response of magnetic nanoparticles. This imaging can be limited in sensitivity by (potentially non-stable) harmonic background, which is generated by e.g. the non-linearity of components (e.g. capacitors), magnetic materials near the components and/or magnetic materials inside the components (e.g. magnetic iron contamination in generally non-magnetic copper). The latter is of concern in cabling, which is used at various positions in an MPI setup. In particular, high currents are conducted by cabling between and in the components of the high-current generator. The strongest magnetic field strengths in the cabling occur where the cabling is wound to serve as an inductor. Besides components of an MPI setup, this effect is largest in the drive field signal generator of the MPI apparatus. 
     SUMMARY OF THE INVENTION 
     It is an object of the present invention to provide a cabling arrangement and a coil apparatus as well as an apparatus and a method for influencing and/or detecting magnetic particles in a field of view that reduce the non-linearity of the cabling leading to less disturbing harmonic background and thus lead to an improvement of the sensitivity of the arrangement. 
     In a first aspect of the present invention a cabling arrangement, in particular for use in a magnetic particle imaging apparatus, is presented comprising: 
     a first AC terminal, 
     a second AC terminal, wherein said first and second AC terminals are configured for coupling an AC voltage between them, 
     a first internal terminal, 
     a second internal terminal, wherein said first and second internal terminals are configured for coupling a DC voltage between them, 
     a first subset of one or more first conductors coupled between the first internal terminal and the second AC terminal, and 
     a second subset of one or more second conductors coupled between the second internal terminal and the second AC terminal, 
     wherein the cabling arrangement enables the superposition of the AC and DC voltages in said first and second conductors, and
 
wherein said first and second conductors are arranged to form a coil to generate a magnetic field in a zone of interest.
 
     In a further aspect of the present invention a coil apparatus, in particular for use in a magnetic particle imaging apparatus, is presented comprising: 
     a cabling arrangement as disclosed herein and 
     a DC voltage or current source coupled, in particular via switches, between the first and second internal terminals. 
     In still a further 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 means 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 means comprising a drive field signal generator unit and drive field coils 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 
     one or more coil apparatuses as disclosed herein forming said drive field coils, wherein the drive field signal generator unit is coupled to the AC terminals of the one or more cabling arrangements of the one or more coil apparatuses. 
     In another aspect of the present invention a corresponding method for influencing and/or detecting magnetic particles in a field of view is presented. 
     Preferred embodiments of the invention are defined in the dependent claims. It shall be understood that the claimed coil apparatus, apparatus and method have similar and/or identical preferred embodiments as the claimed cabling arrangement and as defined in the dependent claims. 
     Said “DC voltage” means a constant voltage with respect to said AC voltage, .i.e. this constant voltage may fluctuate or change but not significantly with respect to the AC voltage, over time. 
     Within MPI, the non-linearity of materials does not (or at least less) become effective if the material is driven into (or nearly into) magnetic saturation. The present invention applies this observation to the cabling, in particular of one or more coils, and proposes to apply dedicated DC currents to bring the magnetic contaminations of the cabling into saturation. In order not to generate an additional static magnetic field (which would shift the first sub-zone, i.e. the field-free point or field-free line), the DC currents are applied such that effectively no AC field, i.e. no far-field, is generated in the zone of interest. The static magnetic field is therefore mainly confined to within the cabling arrangement. 
     The magnitude of the DC current shall preferably be in the order of magnitude of the AC current (e.g. the drive field current) in order to substantially reduce the harmonics generated by the magnetic non-linearities within the cabling. Preferably, a considerable DC current is applied that leads to additional power dissipation in the conductor. 
     Due to the arrangement of the conductors of the first and second subsets the direction of the DC current is such that the conductors of the first subset receive the DC current in a first direction with respect to the area of interest and the conductors of the second subset receive the DC current in a second direction with respect to the area of interest, wherein the first and second directions are preferably opposite to each other. 
     Further, the present invention preferably takes an additional advantage of the DC current in tailoring the generated magnetic field such that the AC fields do not (nearly) cancel out in the entire field-of-view, but only cancel out in the center of the field-of-view, which is the first sub-zone. Hence, the field-free point or line is kept where it was, but additionally a magnetic selection field appears around the field-free point. This selection field can boost (or in some cases also replace) the magnetic selection field as generated by the dedicated selection field signal generator(s). This feature can thus be used to further increase the gradient of the selection field (which translates into better resolution), or to reduce power requirements on the dedicated selection field signal generator(s). 
     In summary, the present invention reduces the magnetic non-linearity of the cabling, used particularly in the drive field signal generator that leads to disturbing harmonic background in the detected signal. In the cabling arrangement, e.g. of the drive field means, different electrically parallel coils/lines that carry said AC currents, in which driven DC currents are additionally supplied (particularly from a DC generator) so as to bring the magnetic non-linearities into saturation. The DC currents applied to a first group of one or more conductors (e.g. coils) have opposite orientation to a second group of conductors (e.g. coils) in order not to generate AC magnetic fields. 
     According to a preferred embodiment the cabling arrangement comprises coupling elements coupled between the first AC terminal and the first internal terminal and between the first AC terminal and the second internal terminal, 
     wherein said coupling elements comprise at least a first capacitor coupled between the first AC terminal and the first internal terminal, and at least a second capacitor coupled between the first AC terminal and the second internal terminal. Said capacitors provide for splitting the first and second internal terminals so that they are at identical AC potential but at different DC potential. 
     According to another embodiment the first AC terminal is coupled to the first and second internal terminal. Hence, the DC voltage is applied in this embodiment either between the first AC terminal and the first internal terminal or between the first AC terminal and the second internal terminal. 
     Preferably, the cabling arrangement further a first capacitive arrangement coupled between the first AC terminal and the first and second internal terminals, and a second capacitive arrangement coupled between the second AC terminal and the first and second sets of conductors. Preferably, a first set of capacitors coupled to the first AC terminal, in particular in series, and/or a second set of capacitors coupled to the second AC terminal, in particular in series. These capacitors are provided to block the DC current from flowing through the first and/or second AC terminals. 
     Optionally, the cabling arrangement further comprises a first inductor coupled to the first internal terminal such that the first inductor and the first subset of conductors are located on either side of the first internal terminal and/or a second inductor coupled to the second internal terminal such that the second inductor and the second subset of conductors are located on either side of the second internal terminal; wherein the cabling arrangement is further configured for coupling said DC voltage to the first inductor and to the second inductor. The DC voltage source is then coupled between the first inductor and the second inductor, possibly via one or several switches. The purpose of this first and second inductor is to separate possible disturbances from the DC voltage source to enter the cabling arrangement, as well as to inhibit the AC voltage from entering the DC voltage source. 
     For the arrangement of the first and second conductors of the first and second subsets different options exist. In one embodiment the first subset comprises two or more first conductors and the second subset comprises two or more second conductors, wherein the first and second conductors are mechanically arranged (i.e. located) substantially in parallel thus forming a cable, wherein said cable is wound as a coil. Further, in one embodiment all first conductors are mechanically arranged (i.e. located) adjacent to each other and all second conductors are mechanically arranged (i.e. located) adjacent to each other, which provides a high saturation. In another embodiment the first conductors and the second conductors are mechanically arranged (i.e. located) alternately, which generally provides an even higher saturation and thus requires a reduced DC current and lower power. 
     Preferably, the cabling arrangement further comprises at least a third internal terminal and at least three coupling terminals coupled between the subsets and the second AC terminal, wherein a first end of the first conductors of the first subset are coupled to different internal terminals and a second end of the first conductors of the first subset are coupled to different coupling terminals, and wherein a first end of the second conductors of the second subset are coupled to different internal terminals and a second end of the second conductors of the second subset are coupled to different coupling terminals. This embodiment provides a higher resistance for the DC voltage. In other words, one or more pairs of conductors are provided, wherein each pair comprises one conductor from the first subset and one conductor from the second subset. Thereby, the two conductors of a pair experience a DC current of opposite direction with respect to the area of interest. 
     Preferably, in this embodiment the cabling arrangement further comprises a series coupling of one or more capacitors coupled between each internal terminal and the first AC terminal and/or a series coupling of one or more capacitors coupled between each coupling terminal and the second AC terminal. These capacitors are provided to block the DC current from flowing through the first and/or second AC terminals. 
     In another embodiment the first subset comprises a single first conductor wound as a first coil and the second subset comprises a single second conductor wound as a second coil, wherein the first coil and the second coil are arranged coaxially. Preferably this can be realised with both coils as solenoids, or with both coils as saddle coils. The coils are thus arranged and provided with DC currents such that the stationary magnetic fields are antiparallel and thus boost the existing magnetic selection field. 
     As mentioned above the coil apparatus may be applied in different kinds of apparatuses, in particular MPI apparatuses. The MPI apparatuses may be differently configured. In one embodiment the apparatus further comprises drive and receiving means comprising said drive field signal generator unit, a signal receiving unit and a drive-receiving coil, said 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, wherein said one or more coil apparatuses form said drive-receiving coil. In other embodiments the proposed idea can be used in an inductive coupling network (or parts of it) used for coupling a generator unit to the respective coils, e.g. for connecting the drive field generator unit with the drive field coils. 
     In still another embodiment the apparatus comprises separate coils for the receiving coil and the drive field coils and/or the apparatus comprises selection-and-focus means including said selection means for generating a magnetic selection-and-focus field having a pattern in space of its magnetic field strength such that the first sub-zone and the second sub-zone are formed in the field of view and for changing the position in space of the field of view within an examination area, said selection-and-focus means comprising at least one set of selection-and-focus field coils and a selection-and-focus field generator unit for generating selection-and-focus field currents to be provided to said at least one set of selection-and-focus field coils for controlling the generation of said magnetic selection-and-focus field. 
    
    
     
       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 circuit diagram of a first embodiment of a coil apparatus according to the present invention, 
         FIG. 7  shows a circuit diagram of a second embodiment of a coil apparatus according to the present invention, 
         FIG. 8  shows a circuit diagram of a third embodiment of a coil apparatus according to the present invention, 
         FIG. 9  shows a circuit diagram of an embodiment of sets of capacitors for used in a coil apparatus according to the present invention, 
         FIG. 10  shows various implementations of the arrangement of the conductors in a cable according to the present invention, 
         FIG. 11  shows a circuit diagram of a fourth embodiment of a coil apparatus according to the present invention, 
         FIG. 12  shows an arrangement of saddle coils using a cabling arrangement according to the present invention, and 
         FIG. 13  shows an arrangement of split solenoid coils using a cabling arrangement according to the present invention. 
     
    
    
     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 x 1 , x 2 , and x 3 , 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&#39;s axis, it will be taken as positive, otherwise as negative. To generate the static selection field, a constant positive current I s  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&#39;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 D   1  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 F   1  is made to flow through both x-coils  12 , a current IF 2  through both y-coils  14 , and a current I F   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   3 ±I S . The current flowing through the remaining two coil pairs  12 ,  14  is I D   k +I F   k , 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 may be used as receive coils as is the case according one embodiment 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&#39;s volume of interest—that volume of the patient (or object) that shall be imaged—is enclosed by the scanner&#39;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 50×10 3  A/m 2 , 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&#39;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, non-surgical 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 μm 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 μm 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 x k -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/μ 0 =2.5 T/m, where μ 0  is the vacuum permeability. The temporal frequency spectrum of the drive field is concentrated in a narrow band around 25 kHz (up to approximately 250 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 of 120 mm/√2≈84 mm. 
     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. 4 a   , 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&#39;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. 4 b    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 means  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 means  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 means are provided in other embodiments. 
     For generating the magnetic drive field the apparatus  100  further comprises drive means  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 means, in particular a receiving coil, and a signal receiving unit  140 , which receives signals detected by said receiving means, are provided. One to three separate receiving coils  124  are provided in an MPI apparatus as receiving means. 
     According to other embodiments of 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, wherein these drive field coils are then called “drive-receiving coils” herein. The generation of magnetic drive fields and the detection of detection signals may then be performed simultaneously or alternately. Preferably, all three drive-receiving coils (or coil pairs) may then act as receiving coils. 
     One to three receiving units  140 —one per receiving coil (or coil pair)—are provided in practice, but more than three 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 receiving coils. 
     Said signal receiving unit  140  comprises a filter unit  142  (also called Rx filter) 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 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  (also called LNA, Low-Noise-Amplifier) 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 means)  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 means  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 means 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. 
     In MPI, the non-linearity of materials does not (or at least less) become effective if the material is driven into (or nearly into) magnetic saturation. The present invention applies this observation to the cabling, used particularly for the drive field coils  124 , and applies dedicated DC currents to bring the magnetic contaminations of the cabling into saturation. In order not to generate an additional static magnetic field, which would shift the field-free point, the DC currents are applied such that effectively no AC (far-) field is generated. The static magnetic field is therefore mainly confined to within the cabling. The DC current shall be in the order of magnitude of the AC current (i.e. drive field) in order to substantially reduce the harmonics generated by the magnetic nonlinearities within the cabling. 
       FIG. 6  shows a circuit diagram of a first embodiment of a coil apparatus  1  according to the present invention. The coil apparatus  1  comprises a first embodiment of a cabling arrangement  1000  and a DC voltage or current source  600  directly coupled between a first internal terminal  301  and a second internal terminal  302 . The cabling arrangement  1000  comprises a first AC terminal  300 , a second AC terminal  310 , the first internal terminal  301  and the second internal terminal  302 . A cabling  1400  (in particular conductors and/or inductors) is connected between the first internal terminal  301 , the second internal terminal  302  and the second AC terminal  310 . The cabling  1400  comprises a first subset of one or more first conductors  401  coupled between the first internal terminal  301  and the second AC terminal  310  and a second subset of one or more second conductors  402  coupled between the second internal terminal  302  and the second AC terminal  310 . Further, the cabling arrangement  1000  is configured for coupling a DC voltage between the first internal terminal  301  and the second internal terminal  302 , in this embodiment via the DC voltage (or current) source  600 . 
     In one exemplary embodiment the cabling  1400  represents the conductors of the drive field coils ( 125 ,  126 , and  127  in  FIG. 5 ), which are preferably—due to the high current—realized by a plurality of parallel conductors, each e.g. consisting of 23000×20 μm Litz wires making up a Rutherford cable. These parallel conductors typically have considerable magnetic coupling. In such an embodiment the first and second conductors  401 ,  402  are mechanically arranged substantially in parallel and forming a cable, wherein said cable is wound as a coil. 
     Due to the high voltage, a first set  500  of capacitors  501  is coupled to the first AC terminal  300 , in particular in series, and a second set  510  of capacitors  511  is coupled to the second AC terminal  310 , in particular in series. To the other ends  320 ,  330  of the first and second sets  500 ,  510  of conductors the drive field signal generator unit ( 122  in  FIG. 5 ) is coupled in case of using the present invention for implementing the drive field coils ( 124 ). 
     In the embodiment of the cabling  1400  presented in  FIG. 6  the first subset of conductors  401  and the second subset of conductors  402  are mechanically essentially in parallel. The AC current (provided by the drive field generator) indicated by arrows  410  flows in parallel trough the conductors  401 ,  402 . A DC current, indicated by arrows  420  additionally flows through them, but with a current direction within the conductors  401  being antiparallel to the current direction in the conductors  402 . This way, the magnetic field essentially cancels out when observed from further away, wherein it is considered that the FOV is essentially in the far field. 
     The DC current flow is a result of a DC voltage applied by a voltage source  600 , which is transparent to AC currents (or is made transparent by additional capacitors connected in parallel). This voltage source  600  can e.g. be a DC charged capacitor or electric battery. The positive side of the battery is connected to the first internal terminal  301  and the negative side of the battery is connected to the second internal terminal  302 . 
       FIG. 7  shows a circuit diagram of a second embodiment of a coil apparatus  2  according to the present invention. The coil apparatus  2  comprises a second embodiment of a cabling arrangement  2000  including a cabling  2400 . The DC current (or voltage) source  600  delivers its current to the internal terminals  301  and  302  via optional inductors  700 ,  701  and switches  800 ,  801 . As in the first embodiment 1, the DC current source  600  might be a battery, an accumulator, a (rechargeable) (super)-capacitor, or a dedicated current source, or parallel connections of these, and is typically connected to further AC circuitry to receive its energy, optionally solely at intervals when no imaging takes places. 
     In order to separate internal terminal  301  from internal terminal  302 , which are at different DC potential, but at identical AC potential, the last capacitor of the capacitors of the first set  520  is split into a first capacitor  502  and a second capacitor  503 . Further, the conductors  401 ,  402  are alternately arranged, i.e. pairs of one first conductor  401  and one second conductor  402  are arranged adjacent to each other. 
       FIG. 8  shows a circuit diagram of a third embodiment of a coil apparatus  3  according to the present invention. The coil apparatus  3  comprises a third embodiment of a cabling arrangement  3000  including a cabling  3400 . This embodiment provides a higher resistance to the voltage source  600 , for an example with 6 parallel conductors  401  to  406 . 
     In this embodiment a third internal terminal  303  and a fourth internal terminal  304  are provided. Further, three coupling terminals  311 ,  312 ,  313  are provided between the subsets of conductors  401 - 406  and the second AC terminal  310 . A first end of the first conductors  401 ,  403 ,  405  of the first subset are coupled to different internal terminals  301 ,  302 ,  303  and a second end of the first conductors  401 ,  403 ,  405  of the first subset are coupled to different coupling terminals  311 ,  312 ,  313 , and a first end of the second conductors  402 ,  404 ,  406  of the second subset are coupled to different internal terminals  302 ,  303 ,  304  and a second end of the second conductors  402 ,  404 ,  406  of the second subset are coupled to different coupling terminals  311 ,  312 ,  313 . 
     The last capacitor of the capacitors of the first set  530  is split into four capacitors  502 ,  503 ,  504 ,  505 , each coupled between the first AC terminal  300  and one of the internal terminals  301 ,  302 ,  303 ,  304 . The second set  540  comprises a first capacitor  511  between the AC terminal  330  and the second AC terminal  310  and three capacitors  512 ,  513 ,  514 , each coupled between the second AC terminal  310  and one of the coupling terminals  311 ,  312 ,  313 . 
       FIG. 9  shows a circuit diagram of an alternative embodiment of the first and second sets of capacitors, in particular for use in the third embodiment of the coil apparatus  3  shown in  FIG. 8 . The first set  550  is similar to the first set  530  shown in  FIG. 8 , wherein each of the capacitors  502 ,  503 ,  504 ,  505  is split into three (alternatively into two or four or more) capacitors. The second set  560  is similar to the second set  540  shown in  FIG. 8 , wherein each of the capacitors  512 ,  513 ,  514  is split into three (alternatively into two or four or more) capacitors. The embodiment provides a more equalized current flow around the peripheries of the capacitor and the cabling elements and hence reduces eddy current losses. 
       FIG. 10  shows various implementations of the cabling, i.e. the arrangement of the conductors in a cable according to the present invention. In particular, the cross sections of the cabling, which consists of parallel conductors that carry AC currents in alike, but DC currents of opposite direction, are shown. The conductor with DC currents flowing into the paper plane are marked by “x” and are termed  402 . The remaining unmarked conductors with currents flowing out of the paper plane are termed  403 . The layout of the conductors shown in the cross section shall be exemplarily for the way the cabling is shaped inside the drive field coil. Various assignments of current direction to the individual positions are possible. 
       FIGS. 10A and 10B  show an arrangement in which alternatingly each second conductor has the same current direction. Due to (possible, but not necessary) rotation of the cable (which could e.g. be formed as a Rutherford cable), the position changes along the length of the cabling. So the cross section of the cabling as shown in  FIG. 10A  will become as shown in  FIG. 10B  at a different position in the cabling. 
       FIGS. 10C to 10F  show arrangements in which the conductors of identical current direction are clustered. The benefit of the clustering might be to have, on average, a stronger saturation, which might be translated to reduced DC currents and hence power requirements. 
     The present invention reduces the magnetic non-linearity of the cabling, used particularly in the drive field coils, that leads to disturbing harmonic background in the detected signal. In preferred embodiments additional advantage is taken of the DC current in tailoring the generated magnetic field such that the AC fields do not (nearly) cancel out in the entire field-of-view but only cancel out in the center of the field-of-view, which is the field-free point. Hence, the field-free point is kept where it was, but additionally a “selection field” appears around the field-free point. This selection field can boost (or theoretically (in some cases) also replace) the selection field as generated by dedicated selection field coils. This feature can be used to further increase the gradient of the selection field (which translates into better resolution), or to reduce power requirements on the dedicated selection field coils. 
       FIG. 11  shows a circuit diagram of a fourth embodiment of a coil apparatus  4  according to the present invention. The coil apparatus  4  comprises a fourth embodiment of a cabling arrangement  4000  including a cabling  4400 . The coil apparatus  4  is rather similar to the second embodiment of the coil apparatus  2  shown in  FIG. 7 , but has just two different conductors  401 ,  402  in parallel. The AC and DC currents have anti-parallel directions in coil  402  and parallel directions in coil  401 . 
       FIG. 12  shows an arrangement of saddle coils using a cabling arrangement according to the present invention, especially as the cabling  4400  of  FIG. 11 . In particular a saddle coil pair of two saddle coils  601 ,  602 , which is used as drive field coils for both the z-(anterior-posterior) and the y-(left-right) direction), is formed, wherein coil  601  is formed by the conductor  401  and coil  602  is formed by the conductor  402 . It is shown how the magnetic fields  430  from the AC current add up in the entire field-of-view. The magnetic fields  440  from the DC currents oppose, i.e. at the very center, where the field-free point resides, no static field is superposed. However, when approaching conductor  402  (top coil  602 ) or conductor  401  (bottom coil  601 ), the magnetic field increases, which is exactly the property of the selection field. 
       FIG. 13  shows an arrangement of split solenoid coils using a cabling arrangement according to the present invention, especially as the cabling  4400  of  FIG. 11 . In particular a solenoid coil of two solenoid coils  603  (formed by conductor  401 ),  604  (formed by conductor  402 ) is shown, which is traditionally realized as one coil and split into two halves here. This coil is used for generating AC fields in the x-direction (head-feet). Since the AC selection field has a fixed orientation, e.g. in the z-(anterior-posterior) direction, as drawn in  FIG. 12 , the selection field generated by the split solenoid coils is different (i.e. orthogonal) in orientation. However, the x-direction (head-feet) selection field from the split solenoid coils is complemented by the (also) different (orthogonal) selection field from the other saddle coil pairs in the y-(left-right)-direction, which in sum generates a selection field in the desired direction (z-, anterior-posterior). 
     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.