Minimizing intravascular magnetic resonance imaging (MRI) guidewire heating with single layer MRI transmit/receive radio frequency coil

A method for controlling an interventional magnetic resonance imaging (iMRI) system configured to control a heating mode of an iMRI guidewire, the method comprising: controlling, during an iMRI procedure, a magnitude of an induced current in a single-layer MRI radio frequency (RF) coil used in the iMRI procedure, or a phase of the induced current by adjusting at least one of: a difference between a working frequency of a whole body coil (WBC) used in the iMRI procedure and a resonant frequency of the single layer MRI RF coil, a coil loss resistance of the single layer MRI RF coil, or a blocking impedance of an LC circuit connected in parallel with the single-layer MRI RF coil; and controlling a heating mode of the guidewire based, at least in part on the magnitude or phase.

BACKGROUND

Conventional approaches to interventional magnetic resonance imaging (iMRI) may result in unwanted radio frequency (RF) heating of endovascular (i.e., intravascular) guidewires or other metal devices used during an iMRI procedure. In a conventional iMRI procedure, a guidewire tip can reach a temperature of up to 74° C. after only thirty seconds of scanning. This level of heating is not safe for patients, limits the amount of imaging data that may be acquired, and thus minimizes the utility of endovascular guidewires used in iMRI. A typical iMRI procedure employs a whole body coil (WBC) as part of an MRI system. A WBC coil is a powerful RF transmitter that may generate more than 25 uT at 1.5 T and consume many KW of power. The guidewire may resonate with the WBC of the MRI system, causing excessive heating of the guidewire.

Conventional approaches to minimizing guidewire heating during iMRI procedures may use different types of guidewire. One approach uses a non-conductive core and non-magnetic wire. A second approach uses a transformer style guidewire in which a longer metal wire is broken into many small sections of metal wire. A third approach uses several layers of conductors and dielectric material. An issue that reduces the effectiveness of all three conventional approaches to mitigating unwanted RF heating is that a new kind of guidewire needs to be developed for each of the three different approaches, which renders the current metal wires available in the market unusable, thereby increasing costs.

Another approach to mitigating unwanted RF heating of guidewires during iMRI is to use parallel transmitting (pTx) coils as transmitting coils. A pTx approach includes many transmit (Tx) modes supported by pTx coils. By controlling the magnitude and phase of each transmit coil, the pTx coils can generate modes that induce little heat on a guidewire tip. A mode in this approach is defined as one combination of all transmitting coils in which each coil has its own unique magnitude and phase. In a pTx approach, a mode having lower induced heat compared to a conventional WBC approach may be defined as a low heat mode. There may be many low heat modes for pTx coils. The low heat modes are not as effective as a conventional WBC Tx field regarding Tx field uniformity but can still generate acceptable MRI images for iMRI applications. Examples of a parallel transmitting approach include M. Etezadi-Amoli et al, “Controlling induced currents in guidewires using parallel transmit”, Proc. Intl. Soc. Mag. Reson. Med. 18 (2010), M. Etezadi-Amoli et al, “Transmit array concepts for improved MRI safety in the presence of long conductors”, Proc. Intl. Soc. Mag. Reson. Med. 17 (2009), and Natalia Gudino et al, “Parallel transmit excitation at 1.5 T based on the minimization of a driving function for device heating”, Proc. Intl. Soc. Mag. Reson. Med. 21 (2013). However, a drawback of the parallel transmitting approach is that it requires a dedicated pTx MRI system. Dedicated pTx MRI systems at 1.5 T and 3 T are not in wide use, and are extremely expensive. Furthermore, institutions may not wish to have duplicative MRI and pTx iMRI systems. This limits the usefulness of conventional pTx approaches for mitigating unwanted RF guidewire heating.

A magnetic resonance imaging (MRI) system, including an iMRI system, may include two kinds of MRI RF coils. The first kind of MRI RF coil is a transmit (Tx) coil. A Tx coil, while operating in Tx mode, transmits high power RF energy into the anatomy of the subject being imaged to excite nuclei spins in the tissue being imaged. The second kind of MRI RF coil is a receive (Rx) coil. An Rx coil, while operating in Rx mode, detects weak signals from nuclei spins of the anatomy being imaged. A conventional MRI system uses a built-in whole body coil (WBC) as a Tx coil. In a conventional MRI system, due to the geometric size of the WBC, the WBC applies RF energy to a much larger region of tissue than is required to image a given region of interest. For example, when a head scan is performed and a WBC is used, not only the head, but also the shoulders and chest also receive a high level of RF energy. This creates a high level specific absorption rate (SAR) issue which limits the clinical utility of MRI systems that use a conventional WBC/Rx coil approach. As a result, a local Tx coil is frequently used to mitigate the SAR problem.

A local Tx coil is designed to apply RF energy into only the anatomy being imaged. There are two conventional approaches to transmitting energy from a power source to a local Tx coil. A first conventional approach is to use a direct connection between the power source to the Tx coil using wires. A direct connection using wires is energy efficient because the energy loss in the connection wires is trivial. A disadvantage of direct connection using wires is that dedicated wiring is required, which increases the cost and complexity of the coil.

A second conventional approach to transmitting energy from a power source to a local Tx coil is to use inductive coupling. For the inductive coupling approach, a primary coil is directly connected to a power source. The primary coil may be a WBC or another large coil. The primary coil is a resonant LC circuit. A smaller second coil (i.e., a local coil) is also used. The second coil is another resonant circuit and is inductively coupled to the primary coil. Thus, energy can be transferred from the primary coil to the second coil. The second coil can be used to excite nearby anatomy more efficiently than the WBC because the second coil is smaller and closer to the nearby anatomy than the WBC. Compared to the first approach using a direct connection with wires, the inductive coupling approach may be less efficient than direct wiring but is still more efficient than a conventional WBC. One benefit of the inductive coupling approach is that no special wiring is required. However, conventional inductive coupling approaches require the use of multiple coils. For example, a conventional inductively coupled knee coil uses two layers of RF coils. The first (inner) layer includes a plurality of Rx coil elements which detect signals from the anatomy while operating in Rx mode, and which are decoupled from the transmitting field while operating in Tx mode. The second (outer) layer is typically a standard birdcage coil that inductively couples to a WBC to create a local amplified transmitting field in Tx mode and which is disabled in Rx mode. However, this conventional inductively coupled dual layer coil has drawbacks. For example, all the individual Rx coil elements in a conventional dual layer coil need associated circuits for decoupling the Rx coil and the local Tx coil while operating in Tx mode. Conventional inductively coupled dual layer coils also require circuits for switching off the Tx coil while operating in Rx mode, which requires complex and expensive control circuitry. This leads to complex and expensive coils. These multiple decoupling circuits and complex control circuits can also decrease the signal to noise ratio (SNR), thereby reducing image quality. Furthermore, the outer layer, by its proximity to the inner layer, will create additional noise when the inner layer is operating in Rx mode. Thus, conventional approaches to mitigating unwanted RF heating of guidewires in iMRI procedures, and the coils used therein, are suboptimal.

DETAILED DESCRIPTION

Embodiments described herein generate an amplified local Tx field using inductive coupling between a WBC and a local single-layer MRI RF coil array in which the magnitude or phase of an induced current in individual coil array elements of the local single-layer MRI RF coil array are independently adjustable to mitigate RF heating of an iMRI guidewire. Embodiments adjust the magnitude and phase of the induced current in individual coil array elements independently to vary the uniformity of the Tx field such that guidewire heating is mitigated during an iMRI procedure while image quality is maintained at a clinically useful level. A discussion of exemplary single layer MRI RF coils and RF coil array elements suitable for use by embodiments described herein now follows.

FIG. 1illustrates an RF coil110and an RF coil120. Coil110and coil120are resonant RF coils and are inductively coupled to each other. Coil110and coil120may be part of an MRI system or iMRI system. In this example, coil110operates as a primary coil and is driven by RF amplifier ε130. Coil120operates as a secondary coil. Secondary coil120is inductively coupled to the primary coil110by mutual inductance M. Secondary coil120is driven by the mutual inductively coupled RF power from the primary coil110. The primary coil110is, in this example, a WBC, while the secondary coil120is a local inductively coupled smaller coil. For the primary coil110, Rp, Cp, and Lpare defined as the coil loss resistance, the equivalent coil breaking point capacitance, and the equivalent total coil inductance respectively. The RF amplifier ε130is defined as an equivalent RF voltage source from an RF power amplifier through a matching circuit, which is not illustrated for clarity. The resistive loss from the RF amplifier ε130is absorbed by Rp, which may be pre-defined for simplicity of calculation. The mutual inductance between primary coil110and secondary coil120is defined as M. The resistive loss, the equivalent breaking point capacitance, and the total inductance of secondary RF coil120are defined as Rs, Cs, and Lsrespectively. The RF current through primary coil110is defined as ip, and the RF current through secondary coil120is defined as is. Mutual inductance between primary RF coil110(e.g. Lp) and secondary RF coil120(e.g. Ls) generates an induced voltage on primary coil110Lp, which may be expressed as +/−jωMis. The sign of jωMisis determined by the polarity between primary coil110and secondary coil120. For clarity of exposition, the positive sign is used in this example. Similarly, the mutual inductance between secondary coil120and primary coil110generates an induced voltage on secondary coil120, which may be expressed as +/−jωMip. The sign is the same for both primary coil110and secondary coil120.

Zp=Rp+j⁡(ω⁢⁢Lp-1ω⁢⁢Cp)⁢⁢and⁢⁢Zs=Rs+j⁡(ω⁢⁢Ls-1ω⁢⁢Cs),
where Z represents impedance.

Equation 1 may be re-written in matrix format, resulting in:

Thus, the solution to equation 2 is:

If both the primary coil110and the secondary coil120resonate at the same frequency, then Zp=Rpand Zs=Rs. Thus

ip=Rs⁢ɛ(Rp⁢Rs+ω2⁢M2)⁢⁢and⁢⁢is=-j⁢⁢ω⁢⁢M⁢⁢ɛ(Rp⁢Rs+ω2⁢M2).
Recall that the phase of isis opposite to the phase of ip, per Lenz's Law. The ratio between isand ipis −jωM/Rs. The secondary coil120's quality factor (Q) may be high, i.e., the value of Rsmay be small. Therefore the ratio between current isand ipmay be large. Furthermore, because secondary coil120is smaller than primary coil110and is also closer to the imaging area than primary coil110, the same magnitude of RF current generates a larger magnetic transmitting field at the area being imaged. Thus, a local inductive coil is significantly more power efficient than a large WBC coil and the local inductive coil's current is dominant compared to the primary coil's current even though their phases are opposite to each other. Furthermore, if a local inductive coil such as secondary coil120does not resonate at the frequency of primary coil110, then the induced current isin secondary coil120can be written as:

As shown in Eq. 4, the frequency deviation of secondary coil120from primary coil110can reduce the magnitude of isand change the phase of is. This approach may be used by embodiments described herein to reduce the local coil120's RF power efficiency if a particular MRI application requires it. For example, as demonstrated by Eq. 4, the induced current isis a function of a coil loss resistance of primary coil110, a coil loss resistance of secondary coil120, or a difference between a working frequency of primary coil110and a resonant frequency of secondary coil120. Thus, embodiments described herein may independently adjust a magnitude of the induced current in a local coil (e.g. secondary coil120), or a phase of the induced current in a local coil (e.g. secondary coil120) by adjusting the coil loss resistance of the primary coil (e.g. primary coil110), the coil loss resistance of the secondary coil (e.g. secondary coil120), or the difference between the working frequency and the resonant frequency. Embodiments described herein may therefore also adjust the magnitude of local Tx field generated by a single layer MRI RF coil operating in Tx mode.

Embodiments described herein include a single-layer MRI RF coil array that employs a single-layer approach to achieve a local inductively coupled Tx transmitter from a plurality of Rx receivers. An example single-layer MRI RF coil array may operate in a Tx mode or an Rx mode. In Rx mode the single-layer MRI RF coil array functions as a plurality of Rx receivers. In Tx mode multiple PIN diodes may be used to re-configure all or less than all the plurality of Rx receivers so that either all or less than all of the plurality of Rx receivers may inductively couple to the WBC and amplify the transmit field. Under this single-layer approach there are multiple approaches that may be employed by embodiments described herein to create a Tx field with a local inductively coupled coil.

A first single-layer approach is to use PIN diodes to configure a plurality of Rx coils together to create a local volume coil, such as a birdcage coil, and to inductively couple the local volume coil to a larger WBC. This first approach may be demonstrated by an example Rx coil that includes 8 loops configured as independent receivers on a cylindrical former. In examples described herein, a loop is considered to be an RF coil element, or RF coil array element.FIG. 2is a diagram of an example 8-loop coil200in Rx mode. 8-loop coil200includes loops201-208. In this example, a loop serves as an independent receiver, and includes receive electronics221-228respectively. Between directly neighboring loops (e.g. loops202,203,204), the directly neighboring loops overlap each other to achieve good isolation, i.e., minimum mutual inductance. Good isolation between loops can also be achieved by using capacitors.

FIG. 3illustrates an Rx coil300that is similar to Rx 8-loop coil200illustrated inFIG. 2, but with additional elements and details. Rx coil300includes multiple PIN diodes306,308, and310. A PIN diode has low impedance (e.g. is shorted) when it is forward biased by a DC power supply. In Tx mode the PIN diodes306,308, and310are forward biased. The circuitry supporting the PIN diodes306,308, and310, such as RF chokes, is not illustrated inFIG. 3for clarity. If all PIN diodes306,308, and310are shorted due to the application of a forward bias, then the circuit of the coil300is changed to the equivalent circuit350. The equivalent circuit350illustrates an 8-rung birdcage coil that will inductively couple to a WBC in Tx mode and amplify the transmitting field and increase the efficiency of the WBC. In summary, the first approach of single-layer technology uses PIN diodes to reconfigure all or some of the Rx coil elements in a plurality of Rx coil elements into a local volume coil to increase WBC efficiency using inductive coupling, and to electrically link Rx coil elements together as one larger inductively coupled Tx coil. This first, conventional approach may be sub-optimal for two reasons. A first reason is that many diodes are required to link different Rx coils together. This increases the complexity of the coils. Therefore, this first, conventional approach may be expensive. The other reason is that even though PIN diodes are considered to be shorted when a forward bias is applied to the PIN diodes, the resistive losses of the PIN diodes are not trivial. A typical value of a forward biased PIN diode is 0.5 Ohm. This 0.5 Ohm could be larger than the coil loss itself for some high Q coils. This additional PIN diode resistive loss reduces the local inductively coupled RF coil's power efficiency.

Example embodiments described herein employ a second, different single-layer approach that uses PIN diodes to facilitate switching an Rx coil element into Tx mode so that all or part of all the Rx elements in a plurality of Rx elements can inductively couple to a WBC individually. In this approach, unlike in the first conventional approach, there are no PIN diodes between Rx elements (e.g., loops, RF coil elements) to link the Rx elements together. When forward-biased, a PIN diode may produce a negligible resistance (e.g., ˜0.1Ω), which is essentially a short-circuit. When reverse-biased, a PIN diode may produce a high resistance (e.g., ˜200 kΩ) in parallel with a low capacitance (e.g., ˜2 pF), which is essentially an open-circuit.

An analysis of the induced current in the Rx elements of an array when operating in Tx mode illustrates the operation of the second approach employed by embodiments described herein. In this analysis it is assumed that the couplings among Rx elements are small and can be ignored. The couplings between Rx elements and the WBC are dominant. For example, the WBC field will induce voltage in one element and generate current flow in that element. That current flow will generate its own field. This additional field will induce voltage on this element's neighbors, including direct or indirect neighbors. This additional induced voltage is ignored in this analysis for clarity of exposition because of the assumption that the couplings among Rx elements are small.

FIG. 4illustrates current distribution through rungs of a 4N-rung birdcage coil in circular polarized (CP) mode. Current distribution for a B field in the horizontal direction (Bx) is illustrated at410. For a B field in the horizontal direction (Bx) the current through a rung k can be written as:

Ikx=I0⁢sin⁡(2⁢⁢π4⁢N⁢k)⁢sin⁡(ω0⁢t)(Eq.⁢5)
where k is the rung number (k=1, . . . , 4N) and ω0is the working frequency.

Current distribution for a B field in the vertical direction (By) is illustrated at420. For a B field in the vertical direction (By) the current through a rung k can be written as

For current distribution for a B field in the vertical direction (By), the time domain function is cosine instead of sine due to the circular polarized mode requirement. Furthermore, the “±” symbol indicates that the direction of Bymay be upward or downward. This affects the rotational direction of the final circular polarized B field, illustrated at430, by making the final circular polarized B field rotate either clockwise or counterclockwise. The total current through a rung k is the sum of the two currents described in equation 5 and equation 6. Therefore:

In this example, a first rung has the same current magnitude I0and angular frequency ω0as another, different rung. The currents in different rungs differ with respect to phase. Thus, a typical high pass, low pass, or bandpass WBC's current distribution in its rungs can be described by Eq. 7.

Consider a 4N loops Rx coil that is put inside this circular polarized uniform B field, where the circular polarized uniform B field is generated by a birdcage coil or WBC.FIG. 4, element430illustrates current distribution for a B field generated from a birdcage coil or WBC in this situation. In this example, good isolation among all loops is assumed. Furthermore, in this example, each loop is identical, the loops use overlap to isolate the directly neighboring elements, all loops are in the same row, and the loops are wrapped around a cylindrical former.FIG. 5illustrates an example 16 loop coil500configured as a birdcage coil.

In transmit mode the CP B1field from a birdcage coil (e.g. a WBC) induces voltage in each loop. The induced voltage for a loop can be written as

Vinduced=d⁡(B1→,Ak→)dt(Eq.⁢8)
whereis the area vector, where the magnitude is the area of the kthloop and the direction is the direction perpendicular to the area towards the outside of the coil.

The B1field and area vectorcan be written as:

Therefore, the current through the equivalent kthrung position is

Iinduced⁢⁢no⁢⁢coupling⁢_⁢k=Vk-Vk-1R=A0⁢B1⁢ω0R*2*Sin⁡(2⁢⁢π8⁢N)*Cos⁡((k-1)*2⁢⁢π4⁢N-ω0⁢t)(Eq.⁢11)
where R is the impedance of a loop. At the resonant frequency the reactive part of the impedance is self-cancelled and only the real part is left. As demonstrated by both equation 11 and equation 7, a uniform circular B1field results. The final B1field Btinside the loops is the sum of both. As a result the final B1field Btinside a small cylinder may be uniform.

In one embodiment, the isolations between coil elements are very small. If the isolations are not small and the mutual inductance is defined as Mkjbetween the kthand jth, elements, high order coupling among elements may be ignored, and thus:

Vkj=Mkj⁢d⁡(Ij)dt(Eq.⁢12)
Therefore, the kthelement will see additional coupled voltage from the jthelement.

Summing all of the coupled voltages of the kthelement results in:

Taking the same approach as illustrated in equation 11, then the induced current at the kth rung position can be written as

As demonstrated by equation 14 above, the second term in the right side of equation 14 still creates a uniform B1field. Therefore, compared to the non-coupled case described in equation 11, the coupled case still creates a uniform B1field. The difference here is that the couplings create the coupled B1field which makes the whole coil array less power efficient than the non-coupled case. However, as long as this uniform coupled field is still more efficient than the primary coil, (i.e., the WBC) the coil elements may still be used to resonate to improve RF power efficiency and reduce SAR.

FIG. 6illustrates an example embodiment of an MRI RF coil element600that may be part of a single-layer MRI RF coil array, and that may be used in an iMRI procedure. MRI RF coil element600employs the second approach to generate a local Tx field. InFIG. 6, the configuration illustrated includes one Rx element only, for clarity. Employing a MRI RF coil element600as part of a single-layer MRI RF coil array using the second approach is simpler than the first approach because fewer PIN diodes are required to switch between Tx mode and Rx mode. Example embodiments thus improve on conventional approaches by saving space within the bore of an MRI apparatus because less hardware is used to construct example embodiments, by avoiding electromagnetic (EM) interference that may be caused by unnecessary hardware, and by reducing manufacturing costs by requiring less hardware, including PIN diodes, than conventional approaches. Example embodiments therefore offer at least one measurable improvement on conventional approaches in addition to providing improved SNR and more uniform fields.

FIG. 6illustrates an MRI RF coil element600configured to operate in a transmit (Tx) mode or in a receive (Rx) mode. In one embodiment, MRI RF coil element600may be part of a single-layer MRI RF coil array configured for use in an iMRI system. A single-layer MRI RF coil array comprises at least one RF coil element (e.g., MRI RF coil element600). The single-layer MRI RF coil array may be arranged in an anatomy-specific shape or configuration, including a closed-shape configuration (e.g. a birdcage coil), or arranged in an open-shape configuration (e.g. in a “C” or “U” shape), or other shaped configuration. In one embodiment, the single-layer MRI RF coil array may be a birdcage coil array. For example,FIG. 5illustrates an example birdcage coil array500. Birdcage coil array500is a sixteen loop, single-row coil array comprising sixteen MRI RF coils. A member of the sixteen MRI RF coils may be, for example, MRI RF coil element600. In another embodiment, an element of the single-layer MRI RF coil array is configured in a saddle-like configuration. In another embodiment, a first element of the single-layer MRI RF coil array is configured in a saddle-like configuration, while a second, different element of the single-layer MRI RF coil array is configured as a loop.

The at least one RF coil element600includes an LC coil610, a matching and Tx/Rx switch circuit620, and a preamplifier630. LC coil610includes at least one inductor640and at least one capacitor650. The at least one inductor640and the at least one capacitor650resonate at a first frequency (i.e., a resonant frequency). The at least one inductor640may be, for example, a co-axial cable, a copper wire, a copper foil soldered to a circuit board, or other conductor.

RF coils for MRI may need to be tuned and matched. Tuning involves establishing or manipulating the capacitance in a coil so that a desired resistance is produced. Matching involves establishing or manipulating the capacitance in a coil so that a desired reactance is achieved. When tuning, the impedance z may be described by Z=R+jX=1/(1/(r+jLω)+jCω). Tuning may be performed to achieve a desired tuning frequency for a coil. ω0identifies the desired tuning frequency. ω0, may be, for example, 63.87 MHz at 1.5 T. The size of a conventional coil facilitates estimating inductance L. With an estimate of L in hand, values for capacitors can be computed to produce a desired resonant peak in an appropriate location with respect to ω0. Once capacitors are selected, the resonant peak can be observed and a more accurate L can be computed. The capacitors can then be adjusted to produce the desired resistance. Once the desired resistance is achieved, then capacitance can be adjusted to cancel reactance.

The matching and Tx/Rx switch circuit620, when operating in Tx mode, electrically isolates the LC coil610from the preamplifier630upon the LC coil610resonating with a primary coil (not illustrated) at a working frequency of the primary coil. The matching and Tx/Rx switch circuit620electrically isolates the LC coil610from the preamplifier630by providing a threshold level of impedance between the LC coil610and the preamplifier630. The primary coil may be, for example, a WBC or other primary coil that is larger than RF coil element600. The LC coil610, upon resonating with the primary coil at the working frequency, generates a local amplified Tx field based on an induced current in the LC coil610. The induced current has a magnitude and a phase. The magnitude of the induced current or the phase of the induced current may be independently adjustable. For example, the induced current is a function of at least a coil loss resistance of the WBC, a coil loss resistance of the LC coil610, or a difference between a working frequency of the WBC and a resonant frequency of the LC coil610. In this embodiment, the frequency of the induced current is the same as the working frequency of the current in the primary coil or WBC, even though the resonant frequency of the LC coil610may be different. Embodiments described herein facilitate adjusting the coil loss resistance of the WBC, the coil loss resistance of the LC coil610, or the difference between the working frequency of the WBC and the resonant frequency of the LC coil610. The magnitude of the induced current or the phase of the induced current are configured to be varied over a range of magnitudes or phases respectively. Example embodiments thus facilitate independently adjusting the magnitude of an induced current in LC coil610, or a phase of the induced current.

Matching and Tx/Rx switch circuit620, when operating in Rx mode, electrically connects the LC coil610with the preamplifier630by providing low impedance between the LC coil610and the preamplifier630. Preamplifier630may be a low input impedance low noise amplifier (LNA). In one embodiment, matching and Tx/Rx switch circuit620may be a capacitive matching and Tx/Rx switch circuit. In another embodiment, matching and Tx/Rx switch circuit620may be an inductive matching and Tx/Rx switch circuit.

Example MRI RF coil elements, MRI RF coil arrays, MRI RF coils, apparatus, and other embodiments, may be configured, for example, as bird cage coils.FIG. 7illustrates one embodiment of a single-layer MRI RF coil array710that includes at least one MRI RF coil elements arranged in a single-row birdcage configuration. The at least one MRI RF coil elements may include, for example, MRI RF coil element600,800,900,1000,1100,1200, or1500.FIG. 7also illustrates an example embodiment of a single-layer MRI RF coil array720that includes at least one example MRI RF coil elements arranged in a two-row configuration. Single-layer MRI RF coil array720includes a first row722aligned with a second row724. First row722includes at least four RF coil elements. Second row724includes at least four RF coil elements.FIG. 7further illustrates another, single-layer MRI RF coil array730. Single layer MRI coil array730is similar to single-layer MRI RF coil array720, except the first row732is not aligned with second row734. For example, first row732may be rotated a number of degrees around a central axis (e.g. z axis) shared with second row734, while734is not rotated, or is rotated a different number of degrees. In different embodiments, first row732may be aligned to within a threshold level of alignment with second row734.

In one embodiment, MRI RF coil array720includes a first plurality of RF coil elements (e.g. first row722) and a second plurality of RF coil elements (e.g. second row724). The first plurality of RF coil elements and the second plurality of RF coil elements are radially disposed about a longitudinal axis702. The first plurality and the second plurality may be longitudinally offset a threshold distance greater than zero along the longitudinal axis702. In one embodiment, an element of the first plurality of RF coil elements is axially offset a threshold amount from a respective element of the second plurality of RF coil elements. For example, an element of the first plurality of RF coil elements may be axially offset 15 degrees, 30 degrees, or another, different number of degrees, from a respective element of the second plurality of RF coil elements. The first plurality and the second plurality may include the same number of RF coil elements, or may include different numbers of RF coil elements. For example, in one embodiment, the first plurality may include eight RF coil elements, while the second plurality may include nine RF coil elements. Other, different numbers of RF coil elements may be employed.

In another embodiment, the at least one RF coil elements is arranged in a three-row configuration. For example, a three-row single layer MRI RF coil array may include a first row that includes at least five RF coil elements, a second row that includes at least five RF coil elements, and a third row that includes at least five RF coil elements. In this embodiment, the first row, second row, and third row may be aligned axially, or may be unaligned axially. In another embodiment, other different numbers of rows, number of RF coil elements, or combinations of alignments may be employed.

For example, in one embodiment, MRI RF coil array720includes a first plurality of RF coil elements722, a second plurality of RF coil elements724, and a third plurality of RF coil elements (not illustrated). In this embodiment, the first plurality of RF coil elements722, the second plurality of RF coil elements724, and the third plurality of RF coil elements are radially disposed about a longitudinal axis. The first plurality722, the second plurality724, and the third plurality are longitudinally offset a threshold amount along the longitudinal axis. In one embodiment, an element of the first plurality722of RF coil elements is axially offset a threshold amount from a respective element of the second plurality724of RF coil elements or the third plurality of RF coil elements. The first plurality722, the second plurality724, and the third plurality may include the same number of RF coil elements, or may include different numbers of RF coil elements. For example, in one embodiment, the first plurality722may include eight RF coil elements, the second plurality724may include nine RF coil elements, and the third plurality may include seven RF coil elements. In another embodiment, the first plurality722, the second plurality724, or the third plurality may include other, different numbers of RF coil elements.

FIG. 8illustrates an MRI RF coil element800. MRI RF coil element800is similar to MRI RF coil element600, but includes additional elements and details. MRI RF coil element800may be part of a single-layer MRI RF coil array. The single-layer MRI RF coil array comprises at least one RF coil element800. In one embodiment, MRI RF coil element800includes a matching and Tx/Rx switch circuit820configured to operate as a capacitive matching and Tx/Rx circuit. In this embodiment, matching and Tx/Rx switch circuit820includes a matching capacitor Cm, a first diode D1, a capacitor Cd, and a first inductor Ld. First diode D1may be a PIN diode. The first diode D1, capacitor Cd, and first inductor Ld create a resonant tank circuit in Tx mode when first diode D1is forward biased. This resonant tank circuit isolates input to the LNA preamplifier630from an induced high current or voltage in LC coil610. The resonant tank circuit further facilitates LC coil610, including capacitors650, inductors640, and matching capacitor Cm, to resonate at a high Q without preamplifier630being electrically connected to the RF coil.

In this embodiment, matching capacitor Cm has a first terminal and a second terminal. Matching capacitor Cm is connected, at a first terminal, to a first terminal of first inductor Ld. First inductor Ld is connected at a first terminal, to a capacitor Cd, at a first terminal. Capacitor Cd is connected, at a second terminal, to first diode D1, at a first terminal. First diode D1is connected, at a second terminal to first inductor Ld, at a second terminal. First inductor Ld is connected, at a second terminal, to a first input terminal of preamplifier630. Preamplifier630is connected, at a second input terminal, to the second terminal of matching capacitor Cm. In Rx mode, first diode D1is backward biased (i.e, first diode D1has a high impedance in Rx mode), so that effectively only Ld is presented between Cm and Preamplifier630. While in this example first inductor Ld, first diode D1, and capacitor Cd are illustrated on a connection path between the first terminal of matching capacitor Cm and a first input terminal of preamplifier630, in another embodiment, first inductor Ld, first diode D1, and capacitor Cd may be connected instead between the second terminal of matching capacitor Cm and the second input terminal of preamplifier630.

In one embodiment, MRI RF coil element800further includes a PIN diode control circuit850. PIN diode control circuit850facilitates selective control of first diode D1. For example, PIN diode control circuit850controls a forward bias applied to first diode D1. PIN diode control circuit850may be operably connected to, for example, first diode D1. In another embodiment, PIN diode control circuit850facilitates selective control of other, different diodes, including shunt diodes, or PIN diodes that comprise a magnitude/phase control component.

FIG. 9illustrates an MRI RF coil element900. MRI RF coil element900is similar to MRI RF coil element800, but includes additional elements and details. MRI RF coil element900includes matching and Tx/Rx switch circuit920. In this embodiment, matching and Tx/Rx switch circuit920includes matching capacitor Cm, first diode D1, first capacitor Cd, and first inductor Ld. Matching and Tx/Rx switch920also includes second diode D9, second capacitor Cd9, and second inductor Ld9. Second diode D9may be a PIN diode. The first diode D1, first capacitor Cd first inductor Ld, and second diode D9, second capacitor Cd9, and second inductor Ld9create a resonant tank circuit in Tx mode when first diode D1or second diode D14is forward biased. This resonant tank circuit isolates input to preamplifier630from an induced high current or voltage in LC coil610. The resonant tank circuit further facilitates LC coil610, including capacitors650, inductors640, and matching capacitor Cm, resonating at a high Q without preamplifier630being electrically connected to the RF coil.

In this embodiment, matching capacitor Cm has a first terminal and a second terminal. Matching capacitor Cm is connected, at a first terminal, to a first terminal of first inductor Ld. First inductor Ld is attached at a first terminal, to first capacitor Cd, at a first terminal. First capacitor Cd is attached, at a second terminal, to first diode D1, at a first terminal. First diode D1is attached, at a second terminal to a second terminal of first inductor Ld. First inductor Ld is connected, at a second terminal, to a first input terminal of preamplifier630. Matching capacitor Cm is connected, at a second terminal, to a first terminal of second inductor Ld9. Second inductor Ld9is connected, at a first terminal, to second capacitor Cd9, at a first terminal. Second capacitor Cd9is connected, at a second terminal, to second diode D9, at a first terminal. Second diode D9is connected, at a second terminal, to a second terminal of second inductor Ld9. Second inductor Ld9is connected, at a second terminal, to a second input terminal of preamplifier630.

In one embodiment, MR RF coil element900further includes a balun930. In this embodiment, balun930is connected, at a first coaxial or two-connection terminal, to a first coaxial or two-connection output terminal of preamplifier630. In another embodiment, balun930is connected between matching and Tx/Rx switch920and preamplifier630. Balun930reduces a common mode current flowing in transmission lines that may connect MRI RF coil element900with an MRI system or iMRI system (not illustrated).

FIG. 10illustrates an MRI RF coil element1000. MRI RF coil element1000is similar to MRI RF coil element800, but includes additional elements and details. MRI RF coil element1000may be part of a single-layer MRI RF coil array. The single-layer MRI RF coil array comprises at least one RF coil element1000. MRI RF coil element1000includes shunt diode D2. Shunt diode D2may be a PIN diode. Shunt diode D2has a first terminal and second terminal. Shunt diode D2is connected, at a first terminal, to the first input terminal of preamplifier630. Shunt diode D2is connected, at a second terminal, to the second input terminal of preamplifier630. To further improve isolation between high induced current in LC coil610and LNA preamplifier630, shunt diode D2provides additional shunt protection for the LNA preamplifier630.

FIG. 11illustrates an MRI RF coil element1100. MRI RF coil element1100is similar to MRI RF coil element600, but includes additional elements and details. MRI RF coil element1100may be part of a single-layer MRI RF coil array. The single-layer MRI RF coil array comprises at least one RF coil element1100. RF coil element1100includes an LC coil1110. LC coil1110is similar to LC coil610but includes a matching inductor Lm having a first terminal and a second terminal. LC coil1110also includes at least one conductor640having a first end connected to the first terminal of the matching inductor Lm, and a second end connected to the second terminal of the matching inductor Lm.

In this embodiment, matching and Tx/Rx switch1120operates as an inductive matching circuit. Matching and Tx/Rx switch1120is connected to matching inductor Lm. Matching and Tx/Rx switch1120includes first inductor Ld having a first terminal and a second terminal, first diode D1having a first terminal and a second terminal, and matching capacitor Cm having a first terminal and a second terminal. Matching inductor Lm is connected at a first terminal with the first terminal of matching capacitor Cm. Matching capacitor Cm is connected at a first terminal with the first terminal of first inductor Ld. First inductor Ld is connected at a second terminal with the first terminal of first diode D1. First diode D1is connected at a second terminal with the second terminal of matching capacitor Cm. Matching capacitor Cm is connected at a second terminal with a first input terminal of pre-amplifier630. Matching inductor Lm is connected, at a second terminal, with a second input terminal of pre-amplifier630. The first diode D1, matching capacitor Cm, and first inductor Ld isolate input to the preamplifier630from an induced high current or voltage in LC circuit1110when first diode D1is forward biased. While in this example first inductor Ld, first diode D1, and matching capacitor Cm are illustrated on a connection path between the first terminal of matching inductor Lm and a first input terminal of preamplifier630, in another embodiment, first inductor Ld, first diode D1, and matching capacitor Cm may be connected instead between the second terminal of matching inductor Lm and the second input terminal of preamplifier630.

FIG. 12illustrates an MRI RF coil element1200. MRI RF coil element1200is similar to MRI RF coil element1100, but includes additional elements and details. MRI RF coil element1200may be part of a single-layer MRI RF coil array. The single-layer MRI RF coil array comprises at least one RF coil element1200. MRI RF coil element1200includes shunt diode D2. Shunt diode D2has a first terminal and second terminal. Shunt diode D2is connected, at a first terminal, to a first input terminal of preamplifier630. Shunt diode D2is connected, at a second terminal, to a second input terminal of preamplifier630. To further improve isolation between high current induced in LC coil1110and LNA preamplifier630, shunt diode D2provides additional shunt protection.

Embodiments described herein may include single-layer MRI RF coil arrays configured in shapes other than the cylindrical shape described above. For example, other shapes, including elliptical, rectangular, square, or other different shapes, may be used to build an Rx coil or single-layer MRI RF coil array for particular applications. For those shapes the concepts of the cylindrical case describe above are still applicable. Non-cylindrical shaped single-layer MRI RF coils may differ from cylindrical single-layer MRI RF coils in that the induced B1field of the other, non-cylindrical shapes is not as uniform as the induced B1field of the cylindrical case, but is still more than uniform enough for a Tx field in a clinical environment, including an iMRI procedure. The non-cylindrical shapes or cross sections discussed above are enclosed shapes or closed-shape configurations. Embodiments are not only applicable to an enclosed shape but may also be implemented as opened shapes, including MRI RF coil elements arranged on a single plane, on two parallel planes, on two planes that are within a threshold of parallel, or MRI RF coil elements arranged in an enclosed shape with a side not present, for example, a “C” shape or a “U” shape.

One embodiment of a single-layer MRI RF coil array that employs an opened shape includes a plurality of loops, saddles, or other MRI RF coil elements arranged on two parallel planes, or on non-parallel planes that are within a threshold tolerance of being parallel to each other, located at least a threshold distance apart, and that face each other directly. A threshold tolerance of being parallel may be, for example, a 1% tolerance, a 10% tolerance, or other, different tolerance. For example, a first point on a first MRI RF coil element may be located 20 cm from a corresponding first point on a facing, second MRI RF coil element, while a second point on the first MRI RF coil element may be located 22 cm from a corresponding second point on the second MRI RF coil element. In this embodiment, the size of the loops may be identical, or may be within a threshold margin of difference. For example, a first loop may describe an area of ×cm2, while a second loop may describe an area of 0.9× cm2. In one embodiment, a plurality of different sized loops may be located on a first plane, while a second plurality of different sized loops may be located on a second, parallel plane, or on a second plane that is within a threshold tolerance of being parallel with the first plane.

FIG. 13illustrates one embodiment of a single layer MRI RF coil array1300that includes at least two RF coil elements. While in this embodiment, only two RF coil loops are illustrated, in another embodiment, other, different numbers of RF coil loops may be employed. The at least two RF coil elements includes a first RF coil element1310and a second RF coil element1320. First RF coil element1310and second RF coil element1320may include a single layer MRI RF coil element, including MRI RF coil element600,800,900,1000,1100,1200, or1500, or other single-layer MRI RF coil elements described herein. First RF coil element1310is arranged on a first plane, while second RF coil element1320is arranged on a second, different plane parallel to the first plane.FIG. 13illustrates an example open shape configuration. The first plane and the second plane may be parallel to each other, and are located at least a threshold distance from each other. The threshold distance is a distance greater than zero. The threshold distance may be a function of a volume to be imaged, for example, the size or shape of a human knee, wrist, or head. In this example, the at least a threshold distance is indicated by “x” inFIG. 13. The first plane and the second plane may, in another embodiment, be within threshold of parallel from each other. The first RF coil element1310and the second RF coil element1320may be offset laterally from each other a distance greater than zero, or may be directly aligned. For example, the first RF coil element1310and second RF coil element1320may be located 30 cm from each other in the x axis, and laterally offset 3 cm in the y axis or z axis. In other embodiments, other offsets may be employed.

First RF coil element1310and second RF coil element1320inductively couple to each other since they face each other. If both first RF coil element1310and second RF coil element1320are tuned independently to the same frequency without the presence of the other coil, their resonant frequency will split into two frequencies: a lower frequency and a higher frequency. The lower frequency is for the current of both RF coil element1310and RF coil element1320flowing in the same direction. The higher frequency is for the current of both RF coil element1310and RF coil element1320flowing in opposite directions. The frequencies may be written as

f=12⁢⁢π⁢(L±M)⁢C,
where L is the inductance of the coil, C is the capacitance, and M is the mutual inductance between RF coil element1310and RF coil element1320.

When both first RF coil element1310and second RF coil element1320are placed inside a WBC and the WBC generates a circular polarized (CP) uniform or a uniform B1field perpendicular to the planes of the coils, then the current induced in one of first RF coil element1310or second RF coil element1320by the WBC directly may be expressed as

i1⁢_⁢WBC=ω0⁢A*B1R(Eq.⁢15)
where A is the area of the loop, B1is the magnitude of WBC field, and R is the coil loss. In an example embodiment in which first RF coil element1310or second RF coil element1320includes RF coil element600,800,900,1000,1100,1200, or1500, then the area A of the loop corresponds to the area of LC coil610or LC coil1110. Here, R is the only term in the denominator of Eq. 15 because

j⁢⁢ω0⁢L-j⁢1ω0⁢C⁢=0
at the resonant frequency ω0. The same current is also true for the other coil. For clarity, herein only the B1field perpendicular to the first plane and second plane is described. However, a B1field that is not perpendicular to the first plane and second plane may be described similarly. This is shown in equation 16 below.

Recall that both currents are flowing in the same direction. Because there is mutual inductance between first RF coil element1310and second RF coil element1320, the final current i1of the first RF coil element1310includes the additional current caused by mutual inductance coupling. The final currents i1and i2can be written as:

Both i1and i2flow in the same direction and have the same current magnitude. In this example, the coupling or mutual inductance between the first loop or first RF coil element1310and the second loop or second RF coil element1320causes a Tx efficiency loss. Thus, the sign before M in equation 17 and equation 18 is “−”, (i.e., negative sign). This embodiment thus may function as the equivalent of a two-turn solenoid or a saddle coil which generates a transmitting field that has a uniformity suitable for clinical use, including iMRI procedures. In another embodiment, other, different configurations of coils may be employed.

In one embodiment of single-layer MRI RF coil array1300, a member of the at least two RF coil elements (e.g. first RF coil element1310, second RF coil element1320) includes an LC coil, a matching and Tx/Rx switch circuit, and a preamplifier. In this embodiment, the LC coil includes at least one inductor and at least one capacitor. The at least one inductor and the at least one capacitor resonate at a first frequency. The LC coil is connected with the matching and transmit Tx/Rx switch circuit. The matching and transmit Tx/Rx switch circuit is connected to the preamplifier. The matching and Tx/Rx switch circuit, when operating in Tx mode, electrically isolates the LC coil from the preamplifier upon the LC coil resonating with a primary coil at the first frequency. The LC coil, upon resonating with the primary coil at the first frequency, generates a local amplified Tx field based on an induced current in the LC coil. A magnitude of the induced current or a phase of the induced current is independently adjustable. For example, the magnitude or phase of the induced current may be variable over a range of magnitudes or phases respectively, by varying the coil loss resistance of the primary coil, the coil loss resistance of first RF coil element1310or second RF coil element1320, or the difference between the working frequency of the primary coil and the resonant frequency of first RF coil element1310or second RF coil element1320. The matching and Tx/Rx switch circuit, when operating in Rx mode, electrically connects the LC coil with the preamplifier. In one embodiment, the matching and Tx/Rx switch circuit is a capacitive matching and Tx/Rx switch circuit. In another embodiment, the matching and Tx/Rx switch circuit is an inductive matching and Tx/Rx switch circuit. In one embodiment, the LC coil includes a shunt PIN diode or protection PIN diode that provides further shunt protection to the preamplifier.

FIG. 14illustrates a single-layer MRI RF coil array1400that is similar to single-layer MRI RF coil array1300but that includes additional details and elements. Single-layer MRI RF coil array1400includes first RF coil element1310, second RF coil element1320, and also includes a third RF coil element1430.FIG. 14illustrates first RF coil element1310, second RF coil element1320, and third RF coil element1430disposed in an open shape configuration. In this embodiment, the RF coil elements1310,1320, and1430of MRI RF coil array1400are arranged approximately in the shape of a “C” or “U”. First loop or RF coil element1310is arranged on a first plane, while second loop or RF coil element1320is arranged on a second, different plane. The first plane and the second plane may be parallel or slightly non-parallel to each other, and are located at least a threshold distance from each other. The threshold distance is a non-zero distance greater than zero. In this example, the at least a threshold distance is indicated by “x” inFIG. 14. In one embodiment, third RF coil element1430is arranged on a third plane that is perpendicular to the first plane and the second plane. In another embodiment, third RF coil element1430is arranged on a third plane that is within a threshold degree of parallel with the first plane or the second plane. For example, in one embodiment configured for a first anatomy to be imaged, the third RF coil element1430is arranged on a third plane that is perpendicular with the first plane and the second plane. In another embodiment configured for a second, different anatomy, the third RF coil element1430is arranged on a third plane that is not perpendicular with the first plane and the second plane. In one embodiment, an angle formed by the intersection of the third plane with the first plane or the second plane is user adjustable.

In one embodiment, third RF coil element1430is offset from the first RF coil element1310or the second RF coil element1320a non-zero amount along a y axis or a z axis. For example, the first RF coil element1310and second RF coil element1320may be located 30 cm from each other in the x axis, and laterally offset 3 cm in the y axis. The third RF coil element1430may be laterally offset 2 cm in the z axis from the first RF coil element1310and the second RF coil element1320. In other embodiments, other offsets may be employed.

Third RF coil element1430, like first RF coil element1310and second RF coil element1320, may include an MRI RF coil element described herein, including MRI RF coil elements600,800,900,1000,1100,1200, or1500. While three RF coil elements are illustrated, in another embodiment, other, different numbers of RF coil elements may be employed.

In one embodiment of single-layer MRI RF array coil1400, a member of the at least three RF coil elements (e.g. first RF coil element1310, second RF coil element1320, third RF coil element1430) includes an LC coil, a matching and transmit (Tx)/receive (Rx) switch circuit, and a preamplifier. In this embodiment, the LC coil includes at least one inductor and at least one capacitor. The at least one inductor and the at least one capacitor resonate at a first frequency. The matching and Tx/Rx switch circuit, when operating in Tx mode, electrically isolates the LC coil from the preamplifier upon the LC coil resonating with a primary coil at the first frequency. The LC coil, upon resonating with the primary coil at the first frequency, generates a local amplified Tx field based on an induced current in the LC coil. A magnitude of the induced current or a phase of the induced current is independently adjustable. The matching and Tx/Rx switch circuit, when operating in Rx mode, electrically connects the LC coil with the preamplifier. In one embodiment, the matching and Tx/Rx switch circuit is a capacitive matching and Tx/Rx switch circuit. In another embodiment, the matching and Tx/Rx switch circuit is an inductive matching and Tx/Rx switch circuit. In one embodiment, the LC coil includes a shunt PIN diode or protection PIN diode that provides further shunt protection to the preamplifier.

Embodiments described herein may also be described using a mode approach. For example, two identical coils facing each other may both resonate at the same frequency if the other coil does not exist. Due to mutual inductance the two coils create two intrinsic resonant modes. The first mode is the lower frequency mode which is called saddle mode or corotation mode, where both coils' currents flow in the same direction. The other mode has a higher frequency and is called anti-saddle mode or counter-rotation mode in which the currents of the coils flow in opposite directions. If a uniform external field or a circular polarized uniform external field is applied to the coils, only the saddle mode configuration will have induced voltage because its net flux is non-zero while the anti-saddle mode's net flux is zero. As a result two identical coils, or coils having sufficient similarity, facing each other will generate an amplified B1 field by the local saddle mode which has a level of uniformity suitable for clinical use, including iMRI use. The external uniform field serves as a selector for modes. The larger the net magnetic flux the mode has, the more energy from the external field is coupled.

This discussion can also be extended to embodiments that employ a plurality of MRI RF coil elements. For example, in an embodiment with N coil elements in which some or all of the N coil elements' isolations may not be good (i.e., the isolation may not be within a threshold tolerance), the N coil elements will couple to each other and create M Eigen-resonant modes in which a mode is a sum of some or all coil elements with different weighting coefficients and phases, where N and M are integers. In this embodiment, a mode is excited proportionally by the net magnetic flux of each mode from the WBC. The most uniform mode among all modes has the largest net magnetic flux from the WBC. For example, a two-element embodiment will be more uniform among modes. Therefore, the most uniform mode among the modes is the strongest mode excited by the WBC. If other less uniform modes' net magnetic fluxes from the WBC are not zero, they will be also excited but the induced fields from them are weaker than the most uniform mode, on average. The other less uniform modes make the final combined induced field more uniform than the induced field from the most uniform mode only. Thus, the final combined induced field is sufficiently uniform for use in clinical MRI applications, including iMRI procedures.

In summary, a plurality of single-layer MRI RF coils or MRI RF coil elements configured as a single-layer MRI RF coil array, resonating with a WBC coil in Tx mode will induce a local amplified Tx field. A coil array element may have the magnitude and phase of an induced current adjusted independently of another, different coil array element. The local amplified Tx field has a threshold level of uniformity and the single-layer MRI RF coil array is used as a transmitter coil. This amplified Tx field improves the WBC power efficiency and reduces the SAR compared to conventional approaches because non-related anatomy areas will not experience a high Tx field from the WBC. Because the Tx field is adjustable via independent phase or magnitude control, the RF heating of an iMRI guidewire that is subjected to the Tx field may be minimized or controlled by varying the uniformity of the local amplified Tx field. In one embodiment, MRI RF array coils that connect to an MRI system can be connected through cables or may be connected wirelessly with no cables.

Embodiments described herein are configured such that the magnitude of the induced current or the phase of the induced current are variable over a range of magnitudes or phases respectively. Example embodiments adjust the magnitude and phase of the induced current isof a coil in Tx mode such that the image quality is still acceptable for MRI or iMRI purposes. When a single-layer MRI RF coil array operates in Rx mode, it operates as a phased array receiving coil with a frequency tuned to the working frequency for optimum SNR. Embodiments described herein provide magnitude and phase adjustment for an RF coil operating in Tx mode. Different approaches to adjust a coil's magnitude and phase are now described herein.

Recall from equation 3 that the local inductively coupled coil current can be written as:

is=-j⁢⁢ω⁢⁢M⁢⁢ɛ(Rp⁢Rs+ω2⁢M2)(Eq.⁢3.1)
where isis the induced current of local RF coil, Rpis the coil loss resistance of the WBC, Rsis the coil loss resistance of the local RF coil, M is the mutual inductance between the WBC and the local RF coil and e is the driving voltage of the WBC. Eq. 3.1 immediately above assumes that both the WBC and the local RF coil resonate at the same working frequency. Because the local coil is smaller than the WBC, the magnetic field generated by isis significantly larger than the magnetic field generated from the WBC. Therefore the Tx field is dominated by the local coil. The WBC is part of the iMRI system and its frequency is fixed at the nominal working frequency. Typically, the working frequency of the WBC cannot be changed, and typically, a WBC cannot be removed from an iMRI system in normal clinical practice.

Embodiments are configured to adjust the magnitude and phase of the induced isof a coil (e.g, a loop, a coil array element) in a single-layer MRI RF coil array while operating in Tx mode, and facilitate ensuring that the coil Tx mode is a low heat mode and that the image quality is still acceptable for an iMRI procedure. Embodiments thus provide the equivalent to a pTx coil without requiring the use of an expensive pTx system. When a coil in a single-layer MRI RF coil array is operating in Rx mode, it operates as a phased array receiving coil where a coil's frequency is tuned to the working frequency for optimum SNR.

FIG. 15illustrates an example single-layer MRI RF coil array element1500configured to operate in a Tx mode and an Rx mode. Single-layer MRI RF coil array element1500is similar to RF coil element600,800,900,1000,1100, or1200, but includes additional elements and details. Single-layer MRI RF coil array element1500includes an LC coil1510, a matching and Tx/Rx switch circuit1520operably connected to the LC coil1510, and a preamplifier or other Rx electronics1530operably connected to the matching and Tx/Rx circuit1520. LC coil1510further includes a magnitude/phase control component1512operably connected to LC coil1510.

LC coil1510includes at least one inductor and at least one capacitor, where the at least one inductor and the at least one capacitor resonate at a first frequency. The first frequency may be, for example, the working frequency. In one embodiment, LC coil1510may be configured as LC coil610illustrated inFIG. 6, 8, 9, or10. In another embodiment, LC coil1510may be configured as LC coil1110as illustrated inFIG. 11 or 12.

In one embodiment, LC coil1510includes at least one conductor. The at least one conductor may be a flexible co-axial cable. Thus, LC coil1510may be configured as a flexible LC coil, and single-layer MRI RF coil array element1500may be configured as a flexible single-layer MRI RF coil array element.

The matching and Tx/Rx switch circuit1520, when operating in Tx mode, electrically isolates LC coil1510from preamplifier1530upon LC coil1510resonating with a primary coil (not illustrated). The primary coil has a working frequency.

In one embodiment, the matching and Tx/Rx switch circuit1520is a capacitive matching and Tx/Rx switch circuit. For example, matching and Tx/Rx switch circuit1520may be configured as matching and Tx/Rx switch circuit820illustrated inFIG. 8. In another embodiment, matching and Tx/Rx switch circuit1520may be configured as matching and Tx/Rx switch circuit920illustrated inFIG. 9. In another embodiment, matching and Tx/Rx switch circuit1520may be configured using other, different matching and Tx/Rx switch circuitry configurations.

LC coil1510, upon resonating with the primary coil, generates a local amplified Tx field based on an induced current in LC coil1510. The induced current is generated by inductive coupling between LC coil1510and the primary coil. A magnitude of the induced current or a phase of the induced current is independently adjustable. The magnitude of the induced current or the phase of the induced current is configured to be varied over a range of magnitudes or phases respectively.

The matching and Tx/Rx switch circuit1520, when operating in Rx mode, electrically connects LC coil1510with preamplifier1530.

The magnitude/phase control component1512is configured to adjust the magnitude of the induced current or the phase of the induced current. In one embodiment, the magnitude/phase control component1512is configured to, upon the single-layer MRI RF coil array element operating in Tx mode, adjust the magnitude of the induced current or the phase of the induced current by shifting the first frequency of the LC coil relative to the working frequency of the primary coil. For example, in one embodiment, magnitude/phase control component1512is configured to use a first approach to adjust a magnitude of the induced current or the phase of the induced current by shifting the MR coil's (e.g., LC coil1510) resonant frequency from the working frequency of the WBC. Recall that the induced current ismay be expressed as:

is=-j⁢⁢ω⁢⁢M⁢⁢ɛ(Rp⁢Rs+ω2⁢M2+jRp⁡(ω⁢⁢Ls-1ω⁢⁢Cs))(Eq.⁢4)
where there is an additional term

jRp⁡(ω⁢⁢Ls-1ω⁢⁢Cs)
in the denominator.

If the coil frequency is same as the working frequency, this additional term has a value of zero. Example embodiments adjust this new term by changing the frequency to either higher than the working frequency or lower than the working frequency. This facilitates adjusting the magnitude and phase of the induced current issimultaneously. If the magnitude and phase of isare changed, the magnitude and phase of the induced field is also changed. Thus we have a coil that is a pTx-equivalent coil. Eq. 4 changes magnitude and phase simultaneously. In some situations, it may be desirable to have additional freedom to control magnitude and phase independently.

In another embodiment, magnitude/phase control component1512is configured to use a second approach to adjust a coil's magnitude and phase by independently adding additional coil loss (i.e., resistance) when operating in Tx mode. For example, in one embodiment, the magnitude/phase control component1512is configured to, upon the single-layer MRI RF coil array element1500operating in Tx mode, adjust the magnitude of the induced current or the phase of the induced current by adding coil loss to the LC coil1510. In this embodiment, the magnitude/phase control component1512comprises a resistor and a PIN diode connected in parallel with the at least one capacitor, where the single-layer MRI RF coil array element1500operates in Tx mode upon the injection of a DC bias into the PIN diode, where the DC bias forward biases the PIN diode.

This second approach includes increasing Rswhen operating in Tx mode. Increasing Rsreduces the magnitude alone if the coil's resonant frequency is the same as the working frequency. Increasing Rsreduces both magnitude and phase if the coil's resonant frequency and the working frequency are different. One example implementation of this second approach is illustrated inFIG. 16, which includes adding coil loss. In this embodiment, the single-layer MRI RF coil array element1600includes a magnitude/phase control component1620, and LC coil1610that includes one or more breaking point capacitors1622. A resistor1624is added in parallel to the capacitor1622through a PIN diode1626. Matching and Rx/Tx1520and1530are not illustrated for clarity. The breaking point can also be a feeding point, i.e., a matching capacitor. In Tx mode the PIN diode1626is shorted and the resistor1624is in parallel with capacitor1622, which causes Rsto increase. The value of Rsmay be chosen to meet particular magnitude and phase requirements for different heating modes, including low heating modes. The resistor1624is configured to be powerful enough to take a high current in Tx mode. For example, resistors having resistance within the kilo-Ohms range may be employed to reduce current magnitude significantly. The smaller the resistor value, the lesser the current magnitude, and the lower the coil Q in Tx mode. In Rx mode the PIN diode1626is open, and the resistor1624is therefore not part of the coil resistance. This still results in a high Q receiving coil.

In another embodiment, the magnitude/phase control component1512is configured to, upon the single-layer MRI RF coil array element1500operating in Tx mode, adjust the magnitude of the induced current or the phase of the induced current by decreasing the induced current. In this embodiment, the magnitude/phase control component1512is configured to decrease the induced current by operating as a parallel resonant circuit when in Tx mode. In this embodiment, the magnitude/phase control component1512comprises an inductor and a PIN diode connected in parallel with a first member of the at least one capacitor, where the first member of the at least one capacitor has a higher capacitance than a second, different member of the at least one capacitor, where the single-layer MRI RF coil array element operates in Tx mode upon the injection of a DC bias into the PIN diode, where the DC bias forward biases the PIN diode. In one embodiment, the magnitude/phase control component1512introduces a blocking impedance of less than one-hundred Ohms to the LC coil when operating in Tx mode.

For example, in one embodiment, magnitude/phase control component1512is configured to use a third approach that includes introducing a parallel resonant circuit in the coil when operating in Tx mode only. One example implementation of this third approach is illustrated inFIG. 17.FIG. 17illustrates a single-layer MRI RF coil array element1700that includes LC coil1710, and magnitude/phase control component1720. Magnitude/phase control component1720is configured as an LC parallel resonant circuit. Magnitude/phase control component1720includes a PIN diode1722that controls magnitude/phase control component1720to minimize the induced current in a coil1710when operating in Tx mode. Typically, the blocking impedance of magnitude/phase control component1720when operating in Tx mode as an LC parallel resonant circuit is very large, e.g., several kOhm. This impedance almost completely eliminates the induced current in coil1710. Example embodiments may still need to use the induced current. Thus, example embodiments may reduce current magnitude, but do not need to almost completely eliminate the current. Therefore example embodiments may employ a weak blocking impedance of several Ohms to tens of Ohms, instead of a strong blocking impedance of several kOhms. Example embodiments may use a capacitor1724having a large value of capacitance and inductor1726having a small value of inductance to resonate. Magnitude/phase control component1720may be installed anywhere in LC coil1710to reduce the induced current in Tx mode only by operating as a parallel resonant circuit. Recalling that

is=-j⁢⁢ω⁢⁢M⁢⁢ɛ(Rp⁢Rs+ω2⁢M2)
and Eq. 4, example embodiments thus facilitate controlling current magnitude by changing the value of Rs. In Rx mode the PIN diode1722is open.

The three approaches to adjust a coil's magnitude and phase described above may be employed together to create a mode selection approach for different induced current magnitudes and phases for the same coil in Tx mode. This approach is shown inFIG. 18.FIG. 18illustrates single-layer MRI RF coil array elements1810,1830, and1850. For clarity of illustration,FIG. 18does not include matching and Tx/Rx switch circuitry or a preamplifier or other Rx electronics.FIG. 18illustrates magnitude/phase control components1812,1832, and1852which are configured to provide two-mode switching. Mode switching is conducted by applying a bias to PIN diode pairs1815and1817,1835and1837, or1853and1855. Each of PIN diode pairs1815and1817,1835and1837, or1853and1855includes a pair of back-to-back PIN diodes. In other embodiments, three or more modes switching may be implemented by adding more back to back diodes into the path. The back to back diodes in PIN diode pairs1815and1817,1835and1837, or1853and1855have reverse polarization from each other. The PIN diodes' breakdown voltage may be greater than the induced voltage across the capacitor1811,1834, or1857respectively when operating in Tx mode. In Tx mode only one pair of diodes (e.g., either PIN diode pair1815or PIN diode pair1817) will be ON (short). For example, PIN diode pair1815can be ON(short) and PIN diode pair1817can be OFF(open) or vice versa in Tx mode. The PIN diode pair1815pair path and the PIN diode pair1817pair path will have different induced magnitude and phase. These different magnitudes and phases may be used for different purposes depending on particular clinical scanning requirements. If one of the PIN diode pairs is not ON while the other one is ON, the breakdown voltage that is greater than the induced voltage across the capacitor will ensure the not-ON PIN diode pair path is not part of the coil and the pair of PIN diodes survives the currents generated when operating in Tx mode. The approach may be extended to more than two modes switching, such as three modes, four modes, or more modes.

Magnitude/phase control component1812is configured to provide control of two different switchable magnitudes and phases by frequency shifting using reactance. Reactance components1811may include a capacitor or an inductor.

Magnitude/phase control component1832is configured to provide control of two different switchable magnitudes and phases using coil losses generated with the resistor-based approach described with respect toFIG. 16. Thus, magnitude/phase control component1832is similar to magnitude/phase control component1812, but includes resistors1831instead of reactance components1811.

Magnitude/phase control component1852is configured to provide control of two different switchable magnitudes and phases control using coil loss produced by the LC parallel resonance approach described with respect toFIG. 17. Magnitude/phase control component1852includes inductors1854configured to operate similarly to inductor1726.

Example embodiments thus provide circuits, components, means, or techniques to facilitate independent magnitude and phase control for an MRI RF coil by solving the magnitude and phase equation with the controllable variables as described above. Furthermore, example embodiments facilitate choosing more than one different magnitude and phase combinations, i.e., more than one different Tx field pattern for different scanning or different heat mode imaging requirements. In embodiments described herein, these approaches create Tx fields in Tx mode. In Rx mode, coils described herein operate as a phased array coil at the nominal working frequency.

In addition to an induced current from a WBC field, an MRI RF coil array element as described herein may experience induced current from other MRI RF coil array elements due to non-trivial mutual inductance among the MRF RF coil array element and the other MRI RF coil array elements.FIG. 19demonstrates this induced current. Two MRI RF coil array elements1910and1920exposed in a uniform B1field of a WBC are shown inFIG. 19. The two coil elements1910and1920have areas A1and A2respectively. The uniform B1field is applied perpendicularly to areas A1and A2. The two coil elements1910and1920have non-trivial mutual inductance between them. The self inductances L1and L2, capacitances C1and C2, coil losses R1and R2, current flow I1and I2, and projected areas A1and A2for coil elements1910and1920respectively are illustrated as shown inFIG. 19.

In one embodiment, assuming that both coil1910and coil1920also resonate at ω0in Tx mode, then Eq. 19(3) and 20(4) can be simplified as below:
jω0B1A1=R1I1+jω0MI2Eq. 21
jω0B1A2=R2I2+jω0MI1Eq. 22

Solving equations 21 and 22 results in:

If mutual inductance is not trivial, i.e., ω0M>>R1and R2, and if the values A1and A2are not significantly different, then Eq. 23 and 24 can be simplified as:

Thus, I1is determined by the WBC induced voltage in coil1920, that is, by its neighbor coil1910instead of coil1920itself, and vice versa. Consequently, the local mutual inductance coupled modes among local array coils are dominant if the mutual inductances are not trivial. Therefore, the final induced field of the array coils (e.g.,1910and1920) from the WBC is the sum of the WBC induced fields from coils with trivial mutual inductance from other elements and the local coupled modes among the coils having non-trivial mutual inductances. The local coupled modes may have multiple different frequencies and different Tx field patterns. Depending on the application in which coil1910and coil1920are being employed, (e.g, for a first anatomy, for a second, different anatomy, for an iMRI procedure) the correct local coupled mode or modes must be adjusted in such a way that its or their frequencies must be same as the working frequency ω0.

FIG. 20illustrates an example embodiment of a four-channel single-layer MRI RF coil array2000configured to operate in a low heat mode. InFIG. 20, single-layer MRI RF coil array2000is illustrated operating in Tx mode. MRI RF coil array2000includes four single-layer MRI RF coil array elements2010,2020,2030, and2040. In this example, the magnitudes and phases of induced current in single-layer MRI RF coil array elements2010,2020,2030, and2040are configured in a low heat mode. Single-layer MRI RF coil array elements2010,2020,2030, and2040may be, for example, embodiments of single-layer MRI RF coil array elements1500described herein. The four single-layer MRI RF coil array elements2010,2020,2030, and2040are arranged in one row and on the same plane. In this example, the row is oriented in the B0direction as illustrated. In another embodiment, other different numbers of single-layer MRI RF coil array elements may be arranged in other configurations of rows.

In this example, single-layer MRI RF coil array elements2010and2040each have areas A1having the same, first, value. Single-layer MRI RF coil array elements2020and2030each have areas A2having the same, second value, where area A1does not equal area A2. In embodiments described herein, the area A1must exceed a threshold level of difference from A2, e.g., be at least 20% different. For clarity of calculation, in this example, all single-layer MRI RF coil array elements2010,2020,2030, and2040have the same coil inductance, defined as L. This coil inductance can be achieved by putting an extra solenoid inductor into a smaller area coil (e.g. single-layer MRI RF coil array elements2020and2030) to boost their inductance.

In this example, for clarity of calculation, all single-layer MRI RF coil array elements2010,2020,2030, and2040have the same resistance R, which can be achieved by adding a small resistor into the single-layer MRI RF coil array elements2010,2020,2030, and2040to make their resistances R equal. Each of the single-layer MRI RF coil array elements'2010,2020,2030, and2040resonances is adjusted by adjusting its capacitance C, illustrated as C1, C2, C3and C4. In this example, single-layer MRI RF coil array elements2010,2020,2030, and2040have the same current magnitude but different phases, such as 180 degrees, 0 degrees, 0 degrees, and 180 degrees per the definition of current flow direction. While in a conventional iMRI system, the phases of the induced current in MRI RF coil array elements may be controlled using complex and expensive pTx techniques to drive the MRI RF coil array elements, embodiments described herein control the magnitude and phase independently using, for example, example single-layer MRI RF coil array elements configured with exemplary magnitude/phase control components or circuitry, without requiring pTx systems.

Single-layer MRI RF coil array elements2010and2020experience non-trivial mutual inductance by choosing an overlap between single-layer MRI RF coil array elements2010and2020that is less than a perfect overlap, i.e., an under-overlap. Single-layer MRI RF coil array elements2030and2040are similarly arranged (i.e., with an under-overlap). Mutual inductance between coil array elements may be considered as trivial when they are either direct neighbors, in which situation the mutual inductance can be minimized using a particular overlap, such as between single-layer MRI RF coil array elements2020and2030, or when they are far enough away from each other that the mutual inductance is small, such as between single-layer MRI RF coil array elements2010and2030, between single-layer MRI RF coil array elements2010and2030, and between single-layer MRI RF coil array elements2020and2040.

Based on the example illustrated inFIG. 20, the following equations for single-layer MRI RF coil array elements2010and2030, which are similar to Eq. 19 and 20, may be derived:

In this example, the signs before jω0M in Eq. 27 and Eq. 28 are negative because single-layer MRI RF coil array element2010's field will decrease single-layer MRI RF coil array element2020's field per the current direction definition in the under-overlap configuration. Letting C1=C2=C, the two coil array element2010and2020create two local modes and their frequencies are:

Putting

In this example, assuming Mω0>>R and A1and A2are different, all R terms may be eliminated when R is added or subtracted to any Mω0term, thus Eq. 30 and Eq. 31 can be simplified as:

In this example, single-layer MRI RF coil array elements2010and2020have the same magnitudes but opposite phases. Since single-layer MRI RF coil array element2030's area is same as that of single-layer MRI RF coil array element2020, and since single-layer MRI RF coil array element2040has an area the same as that of single-layer MRI RF coil array element2010, then the induced current of single-layer MRI RF coil array element2040will, in this example, be same as that of single-layer MRI RF coil array element2010. Similarly, the induced current in single-layer MRI RF coil array element2030will be same as that of single-layer MRI RF coil array element2020. This is the same current magnitude and phase configuration as defined in the low heat mode previously.

Embodiments described herein may be configured as a single-layer MRI RF coil array. One embodiment includes a single-layer MRI RF coil array configured to operate in a Tx mode or in an Rx mode during an iMRI procedure. In this embodiment, the coil array includes a plurality of single-layer MRI RF coil array elements (e.g., single-layer MRI RF coil array elements600,800,900,1000,1100,1200, or1500). A single-layer MRI RF coil array element includes, in this embodiment, an LC coil, a matching and Tx/Rx switch circuit operably connected to the LC coil, a preamplifier operably connected to the matching and Tx/Rx circuit, and a magnitude/phase control component operably connected to the LC coil. In one embodiment, the matching and Tx/Rx switch circuit is a capacitive matching and Tx/Rx switch circuit, while in another embodiment, the matching and Tx/Rx switch circuit is an inductive matching and Tx/Rx switch circuit.

The LC coil includes at least one inductor and at least one capacitor. The at least one inductor and the at least one capacitor resonate at a first frequency.

The matching and Tx/Rx switch circuit, when operating in Tx mode, electrically isolates the LC coil from the preamplifier upon the LC coil resonating with a primary coil. The matching and Tx/Rx switch circuit, when operating in Rx mode, electrically connects the LC coil with the preamplifier. The primary coil has a working frequency.

The LC coil, upon resonating with the primary coil, generates a local amplified Tx field based on an induced current in the LC coil. The induced current is generated by inductive coupling between the LC coil and the primary coil.

A magnitude of the induced current or a phase of the induced current is independently adjustable. The magnitude of the induced current or the phase of the induced current is configured to be varied over a range of magnitudes or phases respectively.

The magnitude/phase control component is configured to adjust the magnitude of the induced current or the phase of the induced current. In one embodiment, the magnitude/phase control component is configured to, upon the single-layer MRI RF coil array element operating in Tx mode, adjust the magnitude of the induced current or the phase of the induced current by shifting the first frequency of the LC coil relative to the working frequency of the primary coil.

In another embodiment, the magnitude/phase control component is configured to, upon the single-layer MRI RF coil array element operating in Tx mode, adjust the magnitude of the induced current or the phase of the induced current by adding coil loss to the LC coil. In this embodiment, the magnitude/phase control component includes a resistor and a PIN diode connected in parallel with the at least one capacitor. The single-layer MRI RF coil array element operates in Tx mode upon the injection of a DC bias into the PIN diode. The DC bias forward biases the PIN diode.

In another embodiment, the magnitude/phase control component is configured to, upon the single-layer MRI RF coil array element operating in Tx mode, adjust the magnitude of the induced current or the phase of the induced current by decreasing the induced current. In this embodiment, the magnitude/phase control component is configured to operate as a parallel resonant circuit when the single-layer MRI RF coil array element operates in Tx mode. In this embodiment, the magnitude/phase control component includes an inductor and a PIN diode connected in parallel with a first member of the at least one capacitor. The first member of the at least one capacitor has a higher capacitance than a second, different member of the at least one capacitor, where the single-layer MRI RF coil array element operates in Tx mode upon the injection of a DC bias into the PIN diode, where the DC bias forward biases the PIN diode. For example, the first member of the at least one capacitor may be a capacitor with a large capacitance value. For example, in one embodiment, 1000 pF is considered a large capacitance value at 63.78 MHz. The impedance of such a capacitor is −j2.5 Ohm at 63.78 MHz. 6.24 nH is +j2.5 Ohm at 63.78 MHz. 6.24 nH is considered, in this example, as a low inductance value. Thus the Q of the MRI RF coil array element can be low. If the Q is 10 at 63.78 MHz, then the LC resonant circuit will provide approximately 2.5/10*10{circumflex over ( )}2=25 Ohm impedance. Providing this level of impedance facilitates decreasing the magnitude of the induced current significantly. While a 1000 pF capacitor is referred to in this example, other values may be employed.

FIG. 21illustrates an example MRI apparatus2100configured with a set of example single-layer MRI RF coils. MRI apparatus2100may be, for example, an iMRI apparatus, or may be operably connected with an iMRI apparatus or iMRI system. The apparatus2100includes a basic field magnet(s)2110and a basic field magnet supply2120. Ideally, the basic field magnets2110would produce a uniform B0 field. However, in practice, the B0 field may not be uniform, and may vary over an object being imaged by the MRI apparatus2100. MRI apparatus2100may include gradient coils2135configured to emit gradient magnetic fields like Gx, Gyand Gz. The gradient coils2135may be controlled, at least in part, by a gradient coils supply2130. In some examples, the timing, strength, and orientation of the gradient magnetic fields may be controlled, and thus selectively adapted during an MRI procedure.

MRI apparatus2100may include a primary coil2165configured to generate RF pulses. The primary coil2165may be a whole body coil. The primary coil2165may be, for example, a birdcage coil. The primary coil2165may be controlled, at least in part, by an RF transmission unit2160. RF transmission unit2160may provide a signal to primary coil2165.

MRI apparatus2100may include a set of RF antennas2150that are configured to inductively couple with primary coil2165and generate RF pulses and to receive resulting magnetic resonance signals from an object to which the RF pulses are directed. In one embodiment, a member of the set of RF antennas2150may be fabricated from flexible coaxial cable. The set of RF antennas2150may be connected with an RF receive unit2164.

The gradient coils supply2130and the RF transmission units2160may be controlled, at least in part, by a control computer2170. The magnetic resonance signals received from the set of RF antennas2150can be employed to generate an image, and thus may be subject to a transformation process like a two dimensional fast Fourier transform (FFT) that generates pixilated image data. The transformation can be performed by an image computer2180or other similar processing device. The image data may then be shown on a display2199. RF Rx Units2164may be connected with control computer2170or image computer2180. WhileFIG. 12illustrates an example MRI apparatus2100that includes various components connected in various ways, it is to be appreciated that other MRI apparatus may include other components connected in other ways.

In one example, MRI apparatus2100may include control computer2170. In one example, a member of the set of RF antennas2150may be individually controllable by the control computer2170. A member of the set of RF antennas2150may be an example MRI RF coil element, or an example single-layer MRI RF coil array. For example, MRI RF coil elements600,800,900,1000,1100,1200, or1500may be implemented as part of RF antennas2150illustrated inFIG. 21. In another embodiment, RF antennas2150may include single-layer MRI RF array2000. In another embodiment, the set of RF antennas2150may include other, different combinations of example embodiments of MRI RF coil elements or example embodiments of single-layer MRF RF coil arrays. The magnitude and phase of an induced current in members of RF antennas2150, including individual coil array elements, may be independently adjustable according to embodiments described herein. Independently adjusting the magnitude and phase of an induced current in members of RF antennas2150faciliates adjusting the value of a Tx field generated by RF antennas2150, and facilitates controlling the heating mode of an iMRI guidewire that may be subjected to the Tx field.

An MRI apparatus may include, among other components, a controller and an RF coil operably connected to the controller. The controller may provide the RF coil with a current, a voltage, or a control signal. The coil may be a whole body coil. The coil may inductively couple with an example MRI RF coil element or single-layer MRI coil array, as described herein, including MRI RF coil element600,700,800,900,1000,1100,1200, or1500, or single-layer MRI coil array2000.

FIG. 22illustrates an interventional magnetic resonance imaging (iMRI) apparatus2200. In one embodiment, iMRI apparatus may be implemented as part of MRI apparatus, including MRI apparatus2100, or may be operably connected to an MRI apparatus, including MRI apparatus2100. iMRI apparatus2200includes a controller2210, a whole body coil (WBC)2220, an intravascular guidewire2230, and a single-layer MRI radio frequency (RF) coil2240operably connected to the controller2210. Controller2210provides the single-layer MRI RF coil2240with a current, a voltage, or a control signal. Controller2210may provide a DC bias to single-layer MRI RF coil2240.

Single-layer MRI RF coil2240includes a plurality of transmit (Tx)/receive (Rx) loops, where a member of the plurality of Tx/Rx loops includes a PIN diode, and an induced current magnitude/phase control component. Single-layer MRI RF coil2240is configured to operate in an Rx mode and in a Tx mode. A member of the plurality of Rx loops may be, for example, single-layer MRI RF coil array element1500, or other embodiment described herein. Single-layer MRI RF coil2240may be an array of single-layer MRI RF coil array elements, including, for example, single-layer MRI RF coil array2000or other embodiment as described herein.

Single-layer MRI RF coil2240operates in the Tx mode upon the injection of a DC bias into the PIN diode. The DC bias forward biases the PIN diode. Single-layer MRI RF coil2240inductively couples with the WBC2220when operating in Tx mode. Single-layer MRI RF coil2240, upon resonating with the WBC2220in Tx mode, induces a local amplified Tx field. The local amplified Tx field is based, at least in part, on an induced current in a member of the plurality of Rx loops, the induced current generated by inductive coupling with the WBC2220. The uniformity of the local amplified Tx field is varied by adjusting the magnitude of the induced current or the phase of the induced current in a member of the plurality of Tx/Rx loops.

The induced current magnitude/phase control component controls a difference between a working frequency of the WBC2220and a resonant frequency of a member of the plurality of Tx/Rx loops, a coil loss resistance of a member of the plurality of Tx/Rx loops, or a blocking impedance of an LC circuit connected in parallel with the member of the plurality of Tx/Rx loops.

The magnitude of the induced current or a phase of the induced current in a member of the plurality of Tx/Rx loops is independently adjustable based, at least in part, on at least one of the difference, the coil loss resistance, or the blocking impedance. The magnitude of the induced current or the phase of the induced current adjusts the uniformity of the local amplified Tx field, and thus controls a heating mode of the guidewire2230. For example, at a first point in time during an iMRI procedure, the guidewire2230may experience a first level of heating while in a first location of the Tx field generated by single-layer MRI RF coil2240, where the Tx field in the first location is dominated by a first member of the plurality of Tx/Rx loops. Controller2210may control the magnitude or phase of a member of the plurality of Tx/Rx loops such that the Tx field is altered, thus changing the level of RF heating induced in the guidewire. At a second, different point in time of the iMRI procedure, the guidewire2230may experience RF heating induced by a different location of the Tx field dominated by a second, different member of the plurality of Rx loops. Controller2210may then control the magnitude or phase of a second, different member of the plurality of Tx/Rx loops such that the Tx field is altered, thus changing the level of RF heating induced in the guidewire at the second point in time. While two members of the plurality of Tx/Rx loops are described here, embodiments may include more than two Tx/Rx loops, and controller2210may control more than two Tx/Rx loops. In one embodiment, guidewire2230may be, for example, a catheter or other surgical device subjectable to RF heating during an interventional magnetic resonance procedure.

FIG. 23illustrates an example method2300for controlling a heating mode of an iMRI guidewire during an iMRI procedure. Method2300includes, at2310controlling, during an iMRI procedure, a magnitude of an induced current in a single-layer MRI radio frequency (RF) coil used in the iMRI procedure, or a phase of the induced current. Method2300may include, at2312, controlling the magnitude or phase of the induced current by adjusting a difference between a working frequency of a whole body coil (WBC) used in the iMRI procedure and a resonant frequency of the single layer MRI RF coil. Method2300may also include, at2314, controlling the magnitude or phase by adjusting a coil loss resistance of the single layer MRI RF coil. Method2300may also include, at2316, controlling the magnitude or phase by adjusting a blocking impedance of an LC circuit connected in parallel with the single-layer MRI RF coil. In another embodiment, method2300controls the heating mode of a catheter, or other device subject to RF heating during an iMRI procedure.

Method2300further includes, at2320, controlling a heating mode of the guidewire based, at least in part on the magnitude or phase of the induced current. Controlling the magnitude or phase of the induced current varies the uniformity of a Tx field generated by the single-layer MRI RF coil. Varying the uniformity of the Tx field changes the heating mode experienced by the guidewire. A heating mode may be a low-heat mode, or other mode.

Circuits, apparatus, elements, MRI RF coils, arrays, and other embodiments described herein are described with reference to the drawings in which like reference numerals are used to refer to like elements throughout, and where the illustrated structures are not necessarily drawn to scale. Embodiments are to cover all modifications, equivalents, and alternatives falling within the scope of the invention. In the figures, the thicknesses of lines, layers and/or regions may be exaggerated for clarity. Nothing in this detailed description (or drawings included herewith) is admitted as prior art.

Like numbers refer to like or similar elements throughout the description of the figures. When an element is referred to as being “connected” to another element, it can be directly connected to the other element or intervening elements may be present. In contrast, when an element is referred to as being “directly connected” to another element, there are no intervening elements present. Other words used to describe the relationship between elements should be interpreted in a like fashion (e.g., “between” versus “directly between,” “adjacent” versus “directly adjacent,” etc.).

In the above description some components may be displayed in multiple figures carrying the same reference signs, but may not be described multiple times in detail. A detailed description of a component may then apply to that component for all its occurrences.

The following includes definitions of selected terms employed herein. The definitions include various examples or forms of components that fall within the scope of a term and that may be used for implementation. The examples are not intended to be limiting. Both singular and plural forms of terms may be within the definitions.

“Circuit”, as used herein, includes but is not limited to hardware, firmware, or combinations of each to perform a function(s) or an action(s), or to cause a function or action from another circuit, logic, method, or system. Circuit may include a software controlled microprocessor, a discrete logic (e.g., ASIC), an analog circuit, a digital circuit, a programmed logic device, a memory device containing instructions, and other physical devices. A circuit may include one or more gates, combinations of gates, or other circuit components. Where multiple logical circuits are described, it may be possible to incorporate the multiple logical circuits into one physical circuit. Similarly, where a single logical circuit is described, it may be possible to distribute that single logical logic between multiple physical circuits.