Electronics – Hadrien A. Dyvorne, Todd Rearick, Hyperfine Operations Inc

Abstract for “Radio-frequency coil signal chains for a low field MRI system”

“Methods, and apparatus are provided for reducing the noise in RF signal chains circuitry for a low field magnetic resonance imaging system. An RF signal circuit may contain at least one field effect transistor. This transistor is designed to work as an RF switch at a frequency lower than 10 MHz. Tuning circuitry may be coupled between inputs to an amplifier and an input to the amplifier. Active feedback circuitry is coupled between the output of the amplifier’s amplifier and an input to the amplifier. The feedback capacitor can reduce the quality factor of the RF coil that is coupled to the amplifier.

Background for “Radio-frequency coil signal chains for a low field MRI system”

“Magnetic resonance imaging, (MRI), is an important imaging method for many applications. It is extensively used in clinical and research settings to create images of the insides of the human body. MRI relies on the detection of magnetic resonance signals (MR), which are electromagnetic waves emitted from atoms when they experience state changes due to applied electromagnetic fields. Nuclear magnetic resonance (NMR), for example, detects MR signals emitted by the nuclei excited atoms during the re-alignment of the nuclear spin of atoms in an object being scanned (e.g., atoms within the human tissue). The processing of MR signals can produce images. This is useful in medical applications. It allows the examination of internal structures and/or biological functions within the body for diagnostic, therapy, and/or research purposes.

“MRI is a popular imaging modality for biological imaging. It can produce images with high resolution and contrast that are non-invasive, without needing to expose subjects to ionizing radiation (e.g. x-rays or introducing radioactive material). MRI can also provide soft tissue contrast which can be used to image subjects that other imaging modalities cannot. MR techniques can also capture information about structures and/or biological process that other modalities cannot. There are a few drawbacks to MRI. These include the high cost of equipment, limited availability, difficulty in obtaining access to scanners for clinical MRI, and/or the lengthening of image acquisition.

The trend in clinical MRI is to increase the field strength to MRI scanners in order to improve scan time, image resolution and image contrast. This, in turn, drives up costs. Most MRI scanners are capable of operating at 1.5 to 3 tesla (T), which is the field strength of B0. An estimate of the cost for a clinical MRI scanner costs approximately one million dollars per tesla. This does not include the significant operation, maintenance, and service costs associated with such MRI scanners.

“Conventional high-field MRI systems require large superconducting magnetic magnets and associated electronics to create a strong uniform magnetic field (B0), in which an object (e.g. a patient) can be imaged. These systems are large, with a typical MRI installation having multiple rooms for the magnet and electronics, as well as the thermal management system and control console areas. MRI systems are expensive and large, so they can only be used by facilities such as hospitals or academic research centers that have the resources and space to buy and maintain them. MRI scanners are not readily available due to their high cost and large space requirements. There are many clinical situations where an MRI scan could be helpful. However, due to the above limitations, it is often not possible or impossible to do so.

“Some embodiments include a switch circuit that can be coupled to an RF coil of a low field magnetic resonance imaging system. The circuit includes at least one field effect transistor, (FET), that can be used as an RF switch at a frequency lower than 10 MHz.

“Some embodiments contain a drive circuit that applies a gate voltage at least to one field-effect transistor, (FET), which is used to operate as a radio frequency switch in a low magnetic resonance imaging system. The drive circuit includes at least one isolation element that isolates a voltage source from at least one FET.

“Some embodiments contain a circuit that tunes a radio frequency coil (RF) coupled to an amplifier of low-field magnetic resonance imaging systems. The circuit includes tuning circuitry that is coupled across inputs to the amplifier, as well as active feedback circuitry that is coupled between the output of an amplifier and an input.

“Some embodiments contain a circuit that tunes a radio frequency coil (RF) coupled to an amplifier of low-field magnetic resonance imaging systems. To reduce the quality factor of the RF coil, the circuit includes active feedback circuitry that is coupled between the output of an amplifier and an input to the amplifier.

“Some embodiments include a technique for tuning a radio frequency coil (RF) coupled to an amplifier of low-field magnetic resonance imaging systems. This involves arranging tuning circuitry at the first and second inputs of an amplifier and coupling active feedback between the output and input of an amplifier.

“Some embodiments include radio-frequency (RF), coils for low-field magnetic resonance imaging systems. The RF coil includes a substrate with a first and second sides, and a conductor that wraps around the substrate at a first plurality locations between the second and first sides. A second portion is wound around the substrate at a second plurality locations between the first and second sides.

“Some embodiments include a method for manufacturing a radio frequency (RF) coil that can be used in a low field magnetic resonance imaging system. This method involves providing a substrate with circumferential holes formed at multiple levels. Connecting grooves are placed at different distances from a first substrate side. Next, wind within a portion of circumferential and connecting grooves a first portion and second portions of conductors from the second substrate side to the first substrate side. The connecting grooves contain a second portion and a third portion of conductors.

“Some embodiments include an RF coil that can be used in low-field magnetic resonance imaging systems. The RF coil is composed of a substrate with a first and second side, and a conductor wrapped around the substrate in a balanced wounding pattern. In this pattern, the conductor wraps around the substrate from one side to another, crossing over the second section of conductor from the second side.

“The above-described apparatus and method embodiments can be used with any combination of features, aspects, and acts as described or further below. These, as well as other aspects, embodiments and features, can be better understood by reading the following description and referring to the accompanying drawings.

High-field systems dominate the MRI scanner market, especially for clinical or medical MRI applications. The general trend in medical imaging is to produce MRI scanners that have greater field strengths. Most clinical MRI scanners operate at 1.5T or 3T, with higher field strengths for research applications. As used herein, ?high-field? MRI systems currently in use in a clinical setting, and more specifically, MRI systems operating with a primary magnetic field (i.e. a B0 field) at 1.5 T or higher, although clinical systems operating between 0.5 and 1.5 T are also referred to as ‘high-field. Field strengths of between 0.2 T to 0.5 T are considered’mid-field. As field strengths in high-field have increased, so have field strengths between 0.2 T and 0.5 T. Field strengths between 0.5 T to 1 T have been described as mid-field. By contrast, ?low-field? Low-field MRI systems have a B0 value less than or equals approximately 0.2T. Systems with a B0 value between 0.2T and approximately 0.3T are sometimes referred to as low-field due to higher field strengths at high frequencies. Low-field MRI systems with a B0 field less than 0.1T are referred herein as “very low-field”. Low-field MRI systems with a B0 field less than 10 mT and lower-field systems are referred herein as ‘ultra-low field.

“Conventional MRI systems require specialized facilities, as we have already discussed. The MRI system must operate in an electromagnetically shielded area. Additionally, the room’s floor must be structurally reinforced. Additional rooms are required for high-power electronics as well as the control area of the scan technician. Access to the site must be secured. A dedicated three-phase electrical connection is required to supply power to the electronics. These electronics are then cooled using chilled water. Additional HVAC capacity is often required. These requirements not only make it costly but also limit the areas where MRI systems are possible to be installed. The operation and maintenance of conventional clinical MRI scanners requires a lot of expertise. The ongoing operational costs of operating an MRI scanner are high due to the need for highly-trained technicians and service engineers. Conventional MRI is expensive and difficult to access. This makes MRI a less accessible diagnostic tool that can deliver a broad range of clinical imaging solutions. MRI is not able to be used in many medical applications. It can assist with diagnosis, surgery and patient monitoring.

“High-field MRI systems, as we have discussed, require special adapted facilities that can accommodate their size, weight and power consumption. A 1.5 T MRI system weighs approximately 4-10 tons, while a 3T MRI system weighs about 8-20 tons. High-field MRI systems require a lot of expensive and heavy shielding. Mid-field scanners can weigh between 10-20 tonnes due to large permanent magnets or yokes. Low-field MRI systems, such as those that operate with a B0 magnetic fields of 0.2 T, are typically 10 to 20 tons in weight. This is due to the high amount of ferromagnetic materials used to generate the B0 field and additional shielding. Rooms must have a minimum of 30-50m2 to accommodate heavy equipment. They also need to have reinforced flooring (e.g. concrete flooring) and special shielding to protect the MRI system from electromagnetic radiation. Available clinical MRI systems cannot be moved and will require a dedicated space in a hospital or other facility. In addition to the high costs involved in preparing the space for operation and the ongoing costs associated with maintaining and operating the system, there will also be additional costs.

“In addition, MRI systems currently in use typically consume large amounts power. Common 1.5 T and 3T MRI systems consume 20-40 kW of power while 0.5 T and 0.2T MRI systems consume between 5-20 kW. Each system uses dedicated and specialized power sources. Power consumption refers to the average power used over a specified time period. The 20-40 kW referenced above is the average power consumed during image acquisition by conventional MRI systems. This may include short periods when peak power consumption is significantly greater than the average power consumption (e.g. when gradient coils or RF coils pulsed for short periods of time). The MRI system’s power storage elements (e.g. capacitors) can be used to address peak or large power consumption intervals. The average power consumption is therefore more important as it determines the type and frequency of the power connection required to operate the device. The available clinical MRI systems need to have dedicated power sources. This means that the MRI system will typically require a dedicated three phase connection to the grid. To convert the three-phase power into single phase power for the MRI system, additional electronics are required. Due to the many physical requirements for deploying conventional clinical MRI system, there is a limited availability of MRI and this severely limits the clinical applications that MRI can be used.

“Accordingly, high-field MRI installations are prohibitive in many cases. They limit their deployment to large institutions or specialized facilities. Also, they restrict their use to strictly scheduled appointments. Patients must visit designated facilities at times and places that have been scheduled in advance. High-field MRI is not fully used as an imaging modality due to the numerous restrictions. Despite these drawbacks, high-field MRI continues to be attractive due to the substantial increase in SNR at higher areas. This drives the industry to use higher field strengths in clinical and medical MRI applications. It also increases the cost and complexity for MRI scanners and limits their availability, preventing them from being used as a general-purpose or generally-available imaging solution.

“The low SNR of MR signal produced in low-field (especially in the very low field) has prevented the development and commercialization of a portable, cost-effective, low-power MRI system. Conventional ‘low-field? MRI systems are referred to as?low-field? because they operate at the highest end of what is commonly referred to as the low field range. (e.g. clinically available low fields have a floor of about 0.2 T) in order to produce useful images. Although they are less expensive than high-field MRI system, conventional low-field MRIs have many of the same problems. Low-field MRI systems, in particular, are heavy, immobile and expensive. They also consume significant power and need dedicated three-phase power hookups. Low-field MRI has made it difficult to develop portable, cost-effective, and/or low-power MRI systems that produce useful images.

“The inventors have created techniques that enable portable, low-field and low power MRI systems. This can increase the widespread deployment of MRI technology in a wide range of environments beyond those currently available at research and hospitals. MRI can now be used in emergency rooms, small clinics and doctor’s offices as well as in mobile units in the field. You can bring MRI equipment to your patient’s bedside to perform many imaging procedures. Some embodiments include low-field MRI systems (e.g. 0.1 T, 50 mT or 20 mT), These systems allow for portable, low-cost and low-power MRI. This greatly increases the availability of MRI in a clinical setting.

“Developing a clinical MRI system for low-field imaging poses many challenges. The term “clinical MRI system” is used herein to refer to an MRI that produces clinically-useful images. This means images with sufficient resolution and adequate acquisition time to be useful for a clinician or physician in order to fulfill a specific imaging application. As such, the resolutions/acquisition times of clinically useful images will depend on the purpose for which the images are being obtained. The relatively low SNR is one of the many challenges that can be faced when trying to obtain clinically useful images from low-field conditions. The relationship between SNR (and B0 field strength) is roughly B0 5/4 for field strengths above 0.2 T, and about B0 3/2 for field strengths below 0.01 T. This means that SNR decreases significantly with decreasing field strength, with even greater drops at low field strengths. The significant drop in SNR due to a decrease in field strength has been a major factor in the development of clinical MRI systems operating in the low-field range. The low SNR at low field strengths has made it difficult to develop a clinical MRI system that can operate in this low-field environment. Clinical MRI systems that aim to operate in lower field strengths have traditionally achieved field strengths of around 0.2 T and higher. These MRI systems can be heavy, expensive, and large. They require dedicated space (or shielded tents), and power sources.

“Inventors have created low-field and very-low-field MRI systems that can produce clinically useful images. This allows for the creation of portable, affordable, and easy-to-use MRI systems not possible with state-of-the-art technology. An MRI system can be carried to the patient in some embodiments to offer a variety of diagnostic, surgical and monitoring procedures.

“FIG. 1. A block diagram showing typical components of a MRI 100 system. FIG. 1. MRI system 100 consists of computing device 104 and controller 106. It also includes pulse sequences store 108, power management 110, magnetics components 120, and power management system 110. System 100 is an illustration. A MRI system could have additional components of any type, in addition or instead of those shown in FIG. 1. These high-level components will be included in an MRI system, but the implementation of those components may vary for each MRI system. We’ll discuss this further below.

“As illustrated at FIG. 1. Magnetics components 120 include B0 magnets 122, shims 124 and RF transmit-receive coils 126, and gradient coils 128. Magnet 122 can be used to create the main magnetic field B0. Magnet 122 can be any combination or type of magnetic components that generates the desired main magnetic field B0. The B0 magnet, which is formed in high-field conditions, is usually made using superconducting material, generally in a solenoid geometrie. This requires cryogenic cooling systems to maintain the superconducting state of the magnet. High-field B0 magnets can be expensive and complicated, consume large amounts of energy, and are therefore costly and complex. Superconducting material is also used to implement conventional low-field B0 magnetics (e.g. B0 magnets operating at 0.2% T). To produce the low field strengths of conventional low-field B0 magnetics (e.g. at 0.2 T), permanent magnets must be used. These magnets need to weigh 5-20 tons to achieve the required field strengths. The B0 magnet in conventional MRI systems is not portable and affordable.

Gradient coils 128 can be arranged in a way that generates gradient fields. For example, they may be arranged in three orthogonal directions (XYYZ) in order to create gradients in B0’s field. Gradient coils 128 can be used to encode emitted MR signal by systematically varying B0 fields (the B0 field generated from magnet 122 or shim coils 124) in order to encode the spatial position of received MR signals as a function frequency or phase. Gradient coils 128 can be set up to change frequency or phase in a linear function. However, nonlinear gradient coils may allow for more complex spatial encoding profiles. A first gradient coil might be designed to selectively alter the B0 fields in a first direction (X), to perform frequency encoding in that direction. A second gradient coil could be set up to selectively change the B0 fields in a second direction (Y), substantially orthogonal, to perform phase encoders. A third gradient coil could be set up to selectively adjust the B0 fields in a third direction (Z), substantially orthogonal, to allow slice selection for volumetric imaging applications. The conventional gradient coils consume considerable power and are typically powered by expensive, large-scale gradient power sources. We will discuss this further below.

“MRI is achieved by stimulating and detecting emitted MR signal using transmit and/or receive coils (often referred as radio frequency (RF), coils). Transmit/receive coils can include separate coils to transmit and receive, multiple coils to transmit and/or receive, or the same coils to transmit and receive. A transmit/receive element may contain one or more coils to transmit, one or several coils to receive, and/or one/more coils to transmitting or receiving. To refer to the different configurations of the transmit/receive magnetics components of an MRI system, transmit/receive coils may also be called Tx/Rx. These terms can be interchanged herein. FIG. FIG. 1. RF transmit and receiver coils 126 are made up of one or more transmit coils. These coils can be used to generate RF pulses that induce an oscillating magnet field B1. You can configure the transmit coils to produce any type of RF pulses.

“Power management system 110” includes electronics that provide power for one or more low-field MRI systems 100 components. As an example, power management system 110 could include power supplies, transmit coil components and/or other power electronics to provide sufficient operating power to energize or operate components of MRI 100. FIG. FIG. The electronics in power supply 112 provide the operating power for magnetic components 120 of MRI system 100. Power supply 112 could include electronics that provide power to B0 coils (e.g. B0 magnet 122) in order to generate the main magnetic field for low-field MRI systems. The transmit/receive switch (116) can be used to control whether RF transmit or RF receive coils will be operated.

“Power component(s), 114 may contain one or several RF receiver (Rx) preamplifiers that amplify MR signal detected by one/more RF transmitter coils. (e.g. coils 126), one (or more) RF transmit power components designed to supply power to one (or more) RF transmit coils. (e.g. coils 126), one (or more) gradient power components which provide power for one (e.g. gradient coils 128) and one (e.

“Power components in conventional MRI systems are expensive, large and use significant power. The power electronics are usually located in a separate room from the MRI scanner. Power electronics require a lot of space. They are complex and expensive devices that require support from wall-mounted racks. The power electronics in conventional MRI systems prevent portability and affordability of MRI.

“As shown in FIG. “As illustrated in FIG. 1, MRI system 100 comprises controller 106, also known as a console. It has control electronics that send and receive instructions from power management system 110. The controller 106 can be configured to execute one or more pulse sequences. These are used to determine instructions to power management system 110 to control magnetic components 120 according to a desired sequence (e.g. parameters to operate RF transmit and receiver coils 126, operating gradient coils 128, etc.). FIG. FIG. Computing device 104, for example, may use received MR data to create one or more MR images by using any image reconstruction process. For data processing by the computing device, controller 106 might provide information about one or several pulse sequences to computing devices 104. Controller 106, for example, may give information to computing device104 about one or more pulse sequences. The computing device may then perform an image reconstruction process, at least partially, based on this information. Computing device 104 is typically a high-performance workstation that can perform complex computationally intensive processing on MR data quickly. These computing devices can be quite expensive.

“As you can see, the current clinical MRI systems (including mid-field, high-field, and low-field) are expensive, large, fixed systems that require large dedicated spaces and dedicated power connections. The inventors developed low-field, and very-low-field MRI systems, which are more cost-effective, less powerful, and/or portable. This greatly increases the accessibility and application of MRI. Some embodiments provide a portable MRI system that can be carried to patients and used at the locations it is most needed.

“Some embodiments of MRI systems are portable. This allows the MRI device, as well as the corresponding power consumption, to be transported to the locations where it is needed. The development of a portable MRI device is not without its challenges. It must be small and light, consume little power, and operate in uncontrolled electromagnetic noise environments (e.g. outside a special shielded area).

“Portability refers to the ability to operate the MRI system in a variety of environments and locations. The current clinical MRI scanners must be installed in specially protected rooms. This is, among other reasons, why they are not portable, non-availability, and cost prohibitive. The MRI system must be able to operate in a wide range of noise environments, and therefore, cannot be used outside of a specially protected room. The inventors have developed noise suppression techniques that allow the MRI system to be operated outside of specially shielded rooms, facilitating both portable/transportable MRI as well as fixed MRI installments that do not require specially shielded rooms. These noise suppression techniques are able to be used outside of specially shielded rooms. However, they can also be used in shielded environments.

The MRI system’s power consumption is another aspect of portability. As mentioned above, clinical MRI systems use a lot of power. They consume between 20 kW and 40 kW per hour. This means that dedicated power connections are required. These connections can be three-phase power connections or dedicated connections to the grid. A dedicated power connection is required to operate an MRI system in other locations than costly dedicated rooms. Low power MRI systems have been developed by the inventors. They can be used with mains electricity, such as a 120V/20A connection in the U.S. or large appliance outlets (e.g. 220-240V/30A), allowing the device’s operation wherever common power outlets are available. You can plug into the wall. facilitates both portable/transportable MRI as well as fixed MRI system installations without requiring special, dedicated power such as a three-phase power connection.”

“A portable MRI device is designed according to the techniques discussed herein and includes RF transmit coils 126 that generate a B1 magnetic fields during a transmit operation and collect flux from an MR signal created by an imaged object during a receieve operation. The RF receive coil amplifies and processes signals before they are converted into MR images. The RF signal chain is the circuitry that controls and processes signals recorded by the coils 126. circuitry. The inventors recognized that the components of the RF signals chain circuitry used in high-field MRI systems are not suitable and/or optimized to be used in low-field MRI systems designed in accordance herewith. Some embodiments provide improved RF signal chains circuitry that can be used in a portable, low-field MRI system.

“FIG. “FIG. The RF signal circuitry 200 comprises RF transmit/receive 210 and transmit/receive 212. These circuitry are used to selectively couple the RF receiver circuitry to the RF coils 210 depending on whether they are being used to transmit or receive. The Larmor frequency is a frequency that RF coils should resonate at in order to perform optimally. According to the following relation, the Larmor frequency (w), is related with the strength of B0 field. The gyromagnetic relationship of the imaged element (e.g. 1H) in MHz/T is the gyromagnetic value of the isotope, while B is the strength the Tesla’s B0 field is in Tesla. For a 1.5T MRI system, the Larmor frequency is approximately 64 MHz and for a 3T MRI system it’s approximately 128 MHz. The Larmor frequency for low-field MRI is significantly lower than that of high-field MRI. The Larmor frequency of a 64 mT MRI is 2.75 MHz. RF signal chain circuitry 200 also includes tuning/matching (214) circuitry that transforms the impedance RF coil 210 to maximize performance. Amplifier 216, which is a low-noise amplifier, receives the output of tuning/matching 214. It amplifies RF signals before they are converted into image signals. One of the problems with low-field MRI systems using RF coils is their susceptibility to noise in electronic parts. According to some embodiments, components 210, 214, and 216 can be configured to reduce noise in RF signal chains.

“Some embodiments contain multiple RF coils in order to increase the signal-to noise ratio (SNR), of signals detected by an RFID coil network. A collection of RF coils can be placed at different orientations and locations to detect an extensive RF field. To improve image acquisition’s SNR, a portable MRI system may include multiple RF transmit/receive coils. A portable MRI system could include 2, 4, 8, 16, 32, or more RF receiver coils to increase the SNR for MR signal detection.

“RF coils can be tuned to increase the frequency at which they are sensitive (e.g. the Larmor frequency). Inductive coupling between neighboring or adjacent coils (e.g., coils located sufficiently close to one another) reduces the sensitivity and dramatically reduces the effectiveness for the collection of RF coils. There are techniques for geometrically decoupling neighboring coils, but these places strict restrictions on the coil position and orientation in space. This reduces the ability to detect the RF field accurately and decreases the signal-to noise performance.

“To reduce the negative effects of inductive co-upling between coils, inventors used coil decoupling techniques to reduce the inductive coupling effect between radio frequency coils in multicoil transmit/receive system. FIG. 3 illustrates an example of a passive decoupling circuit 300 that reduces inductive coupling between radio frequency coils in multi-coil transmit/receive systems. FIG. 3 shows a passive coupling circuit 300 that reduces inductive coupling between radio frequency coils in multi-coil transmit/receive systems. Circuit 300 can be used to decouple RF coils exposed to B1 transmit frequencies (e.g. from an RF transmit loop). The decoupling circuit reduces the current flowing through an RF coil when there is an AC excitation voltage at the Larmor frequency. Inductor L1 is an RF signal coil that is visible to the MRI system’s field of view. To optimize noise performance impedance, capacitors C1 and C2 create a tuning circuit which matches the inductance to the input of low noise amplifier (LNA). To prevent the RF coil from being paired with other coils, the tank circuit which includes capacitor C3 and inductor L2 is reduced by the capacitor C3 and capacitor L1. FIG. FIG. 4A shows a plot of voltage at LNA input at resonant frequency for RF coil. This plot is based on simulation of circuit 300 in FIG. 3. FIG. 4B shows a plot of current through an RF coil, based on simulations of circuit 300 in FIG. 3. The LNA voltage at 2.75 MHz resonant frequency is approximately 26 dB (FIG. 4A and the coil current are?37 dB (FIG. 4B). FIGS. FIGS. 4A and 4B show the magnitude of the measured quantity as a straight line, while the phase of that quantity is shown as a dashed or dashed line.

The inventors recognized that decoupling with a tuned matching filter to lower the current in an RF coil has its drawbacks. For example, it requires tuning multiple components (e.g. capacitors C1,C2 and C3) to the operating frequency. SNR is also affected by losses in L2 inductor. Decoupling efficiency, therefore, is a compromise between SNR efficiency and decoupling efficiency. FIG. FIG. 4B shows that although the tuned match filter reduces coil current substantially at the resonance frequency, the sharp valley of the current waveform indicates that the current reduction through RF coil is very limited for the narrow bandwidth around the resonant frequencies.”

“Some embodiments refer to an improved decoupling system that reduces the current in the RF coil. This is done by dampening the coil response with feedback from the amplifier. FIG. FIG. 5 shows a circuit 400 that provides feedback decoupling according to some embodiments. Circuit 400 has an active feedback path that runs from an amplifier LNA’s output to an input LNA. FIG. 5 shows an example of this active feedback path. FIG. 5 shows the active feedback path. It includes one feedback path. It should be noted that the active feedback path can be implemented in a variety of ways, each providing a different type or feedback decoupling depending on which is selected. In some embodiments, for example, the active feedback path may include a first and second feedback paths that provide feedback signals.

The phase of the feedback signal has an effect on the tuning frequency’s amplification gain. This was recognized by the inventors. In some embodiments, multiple feedback paths are used in the active feedback path. A first feedback path might provide a 90-270 degree out of phase feedback signal with a frequency of the coil, while a second feedback pathway may provide a 180-degree out of phase feedback signal with a frequency of coil. Alternately, the gain of an amplifier can be tuned to be 90 degrees or 270 degrees out-of-phase with the resonant frequency. The maximum amplification gain at the tuning frequencies may be achieved when the phase is 270 degrees. Other embodiments may use a single feedback path. In these cases, the phase of feedback signal can be adjusted to 180 degrees to achieve more efficient decoupling.

Circuit 400 provides feedback decoupling by using active negative feedback to dampen the coil response. Also known as reducing Q factor (or?deQing?). The coil) and reduces the current flowing in it. Circuit 400 includes a tuning/matching loop between the LNA and the RF coil. You can use any suitable tuning/matching device in accordance to certain embodiments. Examples are given below.

“FIG. “FIG. Decoupling circuit 500 uses only one component, i.e. capacitor C1, to tune, in contrast to decoupling Circuit 300. Circuit 500 does not contain reactive components C1 or C2 and therefore does not have an inductor. This prevents the SNR losses that are associated with circuit 300 (inductor L1 being in circuit 300).

“Capacitor C1 can be implemented with a capacitor of fixed value. Alternately, capacitor C1 can be implemented with a capacitor having a fixed value (e.g., via a varactor diode). Another way to implement capacitor C1 is to use a capacitor of fixed value (e.g. 300 pF) in parallel with a capacitor of variable value. This arrangement reduces AC losses caused by the use of a variable capacitor within the feedback loop.

“FIG. 7A shows a plot of voltage at LNA input at resonant frequency RF coil. This is based on simulations of circuit 500 in FIG. 5. FIG. FIG. 7B shows a plot of current through RF coil, based on simulation of circuit 500 in FIG. 6. The LNA input voltage at 2.75 MHz resonant frequency is approximately 8 dB. (FIG. 7B, and the current through the coil at?35 dB. Contrary to FIG. FIG. 7B shows that coil current is lower when decoupling circuit 500 is used than circuit 300. Circuit 500, by contrast to circuit 300, provides RF coil-decoupling over a larger bandwidth.

“FIG. “FIG.8” illustrates an alternate feedback-based decoupling circuit 600. In this circuit, the single capacitor tuning/matching loop of circuit 500 is replaced by an alternative feedback circuit. 6 is replaced by a tuning/matching system that includes components C1,C3 and L2. Circuit 600 uses a tuning/matching system to tune the RF coil (represented by L1), in addition to feedback-based decoupling via an active feedback path that includes capacitor C2.

“In certain embodiments, capacitive feedback circuitry is provided, for instance, by the feedback parts of circuits 400 to 500 and 600 in FIGS. 5 and 8 are replaced by mutual inductive feedback circuitry. FIG. FIG. 24 shows an alternative feedback-based circuit 2400. In this circuit, the capacitive feedback circuitry of circuit 500 is replaced by one that uses feedback. 6 is replaced by mutual inductive feedback circuitry, which includes components R1, L2 and R2. Circuit 2400 has inductors L1 & L2 that are connected mutually using a transformer, or by the air.

A transmit/receive switch is another technique to achieve RF coil decoupling according to some embodiments. When RF signals are being transmitted from one or more RF transmit loops, the transmit/receive toggle switch isolates the RF coil and the amplifier. The transmit/receive toggle divides the tuning/matching networks into two sections to protect sensitive electronics during RF transmit cycle. The transmit/receive button 312 in some conventional MRI systems, such as high-field MRI system, is usually implemented with a diode like a PIN diode. FIG. 9 shows a diagram of transmit/receive circuitry with a diode, D1. 9 is circuit 700. The transmit pulse causes diode D1 to be turned on. This creates a short circuit that isolates the RF signal coil and the receive electronics. The resulting network, as described in connection to circuit 300, creates a tank circuit that has a high impedance. This ensures that the current in RF coil stays small. The RF coil is connected to the amplifier during receive cycles. This allows the tank circuit to tune the RF coil to reduce the current flowing through it while still allowing sufficient signal to reach the amplifier’s output. The RF coil is connected during transmit cycles to a first tank circuit and a second tank loop during receive cycles.

“Conventional decoupling systems, such as the one shown in FIG. PIN diodes are often used to isolate the receiver electronics from the RF signal loop in conventional decoupling circuits, such as the one shown in FIG. To turn on a PIN diode in a decoupling circuit, however, it requires approximately 0.1 A. For example, a transmit/receive system with eight coils might require 0.8 A to decouple each receive coil from the RF signal coil for each transmit or receive cycle in an image acquisition pulse sequence. The decoupling circuits in the RF transmit/receive systems consume significant power over the course of an image acquisition protocol. When PIN diodes have been used, a biasing resistor (R1) and an AC blocking filter consisting of components L1 & C1 are required. The ground circuit is not isolated when it is in its off state. While PIN diodes are able to work at higher frequencies in high-field MRI systems (e.g. 10 MHz), they do not perform well at lower operating frequencies, such as those used in very low field or low-field MRI systems. The PIN diode works by rectifying the signal, not blocking it at such low frequencies. A DC bias current Ibias, for example, allows the diode’s forward bias to continue even when a negative signal has been applied. To block an AC signal with frequency f and peak current Ipeak the ratio Ipeak/f must be lower than the sum of the DC bias current Ibias, and the carrier life t. Some low-field MRI applications might have these parameters: Ipeak=10 A. f?2.75 MHz. Ibias=100 mA. According to the relation above, for these parameters, the PIN diode would need to have a carrier lifetime ?>37 ?s, which is not a characteristics of commercially-available PIN diodes.”

“Inventors recognized that PIN Diodes commonly used in a decoupling system may be replaced with Gallium Nitride GaN field effect transistors (FETs). This addresses some of the limitations of using PIN DIodes in RF transmit/receive systems of low-field MRI systems, including reducing power consumption. GaN FETs are much more efficient than PIN diodes, as they require only a few microamps to switch on. This reduces the power consumption by many orders of magnitude. The resistance of the GaN FETs is smaller than PIN diodes when they are turned on, which reduces the tank circuit’s impact. In some embodiments, diode 700 D1 is replaced by one or more GaN FETs to reduce the power consumption of RF transmit/receive systems.

“FIG. “FIG. Circuit 412 is composed of two mirrored FETs. However, some embodiments allow for an RF transmit/receive circuit 412 to include any number of FETs. GaN FETs are more efficient than PIN diodes. They operate at all frequencies and consume very little power.

“FIGS. 11A-C show operating states of FETs used as switches in RF transmit/receive systems in accordance to some embodiments. FIG. FIG. 11A shows a GaN FET that is configured as a switch between two drain nodes D and source nodes S. The gate of the GaN FET controls the state of the switch from on to off. FIG. FIG. 11.B shows that the GaN FET can also be modelled with three lumped capacitors: C_dss, C_gss and C_gd in the off state. The drain D can be isolated from the source S in such a configuration provided that C_ds is low (e.g. 10-100 pF). Some embodiments have a drain-source capacitance greater than 15 pF. FIG. FIG. 11.C shows that the drain-source capacitance (C_ds) is replaced in the on state by a short circuit.

“FIG. “FIG. 12” illustrates a circuit 1000 that drives a gate voltage on GaN transistors U1 or U2 and is arranged to act as an RF transmit/receive switching in accordance with certain embodiments. GaN FETs can be used to decouple and couple the receive electronics to the RF coil. Inductors L5 & L6 can be used as transformers to connect a control signal V2 with the gates of the FETs U1 & U2, and provide ground isolation. The diode (D3) rectifies the control signal to create DC on/off voltage across the capacitor C7. The resistor R11 can be used to discharge capacitor C7 as well as the gate capacitance for the FETs. The transmit/receive switch will turn off or on quickly depending on the time constants of C7+Cgates and R11. In certain embodiments, V2 can be a 10MHz sine wave that is coupled to L5 to drive FETs. The 10 MHz signal can be used to turn on/off the FET gates, and then switched off. To open the switch, R11 is used to discharge the gate drive resistor. FIG. 12 illustrates how the coupling between L5/L6 and inductors L5/L6 can be poor. Inductances may also be small. In some cases, L5/L6 could be used as a small air-core transformer (or as an RF transformer).

“FIG. “FIG. 13” illustrates a circuit 1100 that drives a gate voltage on GaN transistors U1 or U2. It is arranged to function as an RF transmit/receive switching in accordance with certain embodiments. Circuit 1100 uses the RF transmit pulse as the control signal to activate the transmit/receive switch, rather than an externally supplied control signal V2 like circuit 1000. FIG. FIG. 13 shows a coil that is represented by an inductor L6. This coil is designed to receive the RF transmitter pulse and generate a voltage which drives the GaN FETs. One embodiment may associate each of the RF coils with a coil L6 that receives the RF transmit pulse. Another embodiment may associate a subset of RF coils with a coil L6 that is configured to receive an RF transmit pulse. The switch signal generated in response by coil L6 may then be distributed to other circuitry in the array. Circuit 1000 does not require a separate control signal generator, so circuit 1100 has fewer complex circuitry. The transmit pulse is used as a control signal. However, the switch doesn’t close until RF transmission starts. Pulse receiver coil L6 can detect the RF transmit pulse.

“Some embodiments relate to a novel design of a radio frequency (RF) coil that can be used in a low field MRI system. Conventional RF coil designs are designed to work in MRI systems as a solenoid. This wraps around the object and creates a helix pattern. Head coils are used in MRI systems because they have a conductor that is in a solenoid arrangement. This allows for a head to be inserted into the solenoid. FIG. FIG. 14A shows a schematic illustration of a solenoid-RF coil design, in which a conductor wraps in a series of loops around the substrate in one pass. The loops are from the substrate’s first side to the second side. The conductor can be returned to its original side when the substrate’s second side is reached. FIG. FIG. 14B shows a top-view of the coil arrangement in FIG. 14A shows the conductor loops as vertical lines. Points V+ & V? V+ and V? are the ends of the conductors in the coil. They are connected to an amplifier in an MRI system (e.g. a low-noise amplifier) to amplify the recorded signals.

“In ideal cases, the potentials recorded at the outputs the RF coil are balanced so that V+/V?=0 in absence of electromotive force(emf). If an object such as the head or body of a person is inserted into the solenoid, parasitic coupling may occur between the object and conductor in the coil. This could lead to V+ and V? Unbalanced, resulting in a voltage at amplifier input. When the coil is used in an MRI system, the voltage is expressed as a noise signal in MR signal. The parasitic coupling can affect signals at V+ and V depending on where the head is located within the RF coil. differently. If the object is placed at one end, the amount of noise introduced by parasitic coupling into the recorded signal may be greater at V+ than at V. Because of the shorter conductor distance that separates V+ from the point at which noise was introduced into the coil, Alternately, if the object is placed at or near the centre of the coil, between V+, V?, then the noise introduced into the coil will affect both V+, V?. equally. Another implementation would see more noise at V if the object were placed closer to V?. ”

“FIG. 15A shows schematically that an object (represented as voltage source) is placed into a solenoid coil in a specific location. A parasitic coupling (represented as impedanceZC) is then introduced into the coil at one point. In practice, the parasitic coupling between the object and the coil winding will not be distributed. FIG. FIG. 15B shows an impedance model showing how parasitic coupling can affect the voltages V+ and V? The conductor’s ends. ZC is the parasitic coupling of the object and coil, Z+ the impedance within the conductor between where the parasitic coupling was introduced and Z+, Z? ZC represents the impedance of the conductor between V+, Z? and the point where the parasitic coupling was introduced. ZG represents each end of the conductor (i.e. V+ and V?). Ground. If there is weak parasitic coupling between an object and the coil (e.g. Zc?. Z+. Z?, ZG), then the following relation describes the difference of potential at the two ends V+ and V? :”

“V + V – =Z G Z C 2? ( Z ? – Z+ ) V 0

“Because V+ and V are the outputs of the coil? Some RF coils have a balun between their RF coil and amplifier in order to balance the output and reject common mode noise. Baluns are not recommended for low-field MRI systems because of the small magnitude signals received by the coil as well as the lossy characteristics that baluns can have. Some embodiments address an RF coil design which uses a winding pattern to reduce common mode noise. This eliminates the need for a balun.

“FIGS. 16-19 illustrate schematically RF coil designs according to some embodiments. FIGS. The result is a solenoid coil with similar magnetic properties to the coil designs shown in FIGS. 14A. 14A. For example, the magnetic flux detected by turns of the coil placed close together is similar. The electrical properties of the RF coil designs in FIGS. 16-19 are different. The electrical properties of the RF coil designs shown in FIGS. 16-19 are not the same as those shown in FIG. 14A coil design. Particularly, FIGS. 16-19 show the winding designs. FIGS. 16-19 show winding designs that improve balance and reject common mode when images are inserted into the coil. The parasitic coupling between an object and the conductor causes a voltage to be induced near the object when the object is inserted into a coil. FIGS. 16-19 show the winding patterns. FIGS. 16-19 show that adjacent turns of the conductor have the same inductance/potential when a voltage to them is applied. This is because they are at a similar distance from V+ and V??. The voltage caused by parasitic coupling an object in the coil to the conductor creates a noise signal similar to V+ and V?, regardless of its location in the coil. Common mode noise is reduced by?0

Instead of winding the conductor in a single loop from one end to the other, as shown in FIG. 14A: In some embodiments, the conductor can be wound in a balanced manner using multiple passes of loops (e.g. two or more) from one end to the other. FIG. FIG. 16A illustrates an “interlaced” arrangement. 16A shows an “interlaced” winding pattern. A conductor is wound around a substrate starting at one end. The conductor then moves along the winding direction in a single pass, skipping different portions of the substrate at different levels. The conductor is wound around the substrate portions that were missed in the first pass when it is wound from the second (other) end of the substrate. FIG. FIG. 16B is a top view showing the interlaced winding pattern shown in FIG. 16A.”

“FIG. 18A shows an alternate balanced winding pattern that conforms to some embodiments. FIG. FIG. 18A shows a first plurality loops of conductor that are wound around the substrate, passing from the first to the second ends without skipping any levels. To create a “double?”, a second plurality loops near the first plurality are wound around substrate on the second pass. Winding pattern. FIG. FIG. 18B shows a top-view of the winding pattern shown in FIG. 18A.”

“FIG. 19 is a top view showing another balanced winding configuration with an interlaced configuration, in accordance to some embodiments. FIG. FIG. 19 shows a winding pattern that forms loops at a series of levels, rather than from one end of a substrate to the other (e.g. as shown in FIG. 16A) The conductor is wound in two passes from the first to the second ends of the substrate. A first pass forms loops at a series of levels from one end to another end of the substrate (e.g., as shown in FIG. The invention does not limit the use of a particular angle for winding the helix. Any angle that has a desired number of turns around a substrate can be used.

“FIG. 21 shows a process 2100 to manufacture an RF coil according to some embodiments. Act 2110 provides a substrate around which the conductor is to wind. Any suitable non-magnetic material may be used as the substrate. The substrate may be made of a plastic material that has been fabricated using additive manufacturing processes (e.g. 3D printing). The process 2130 proceeds to act 2112 where a plurality grooves are created in the substrate. The substrate may have a top and bottom, and the plurality grooves can be placed at different locations from the top to the bottom. The substrate may be shaped like a helmet, where the head of an individual can be placed. In some embodiments, the grooves are made as a plurality circumferential grooves. From the top to bottom, the circumference of helmet. Some embodiments of the plurality rings are separated by the same spacing from top to bottom of substrate to create multiple levels in which a conductor can be wound. A plurality connecting grooves can also be included between the circumferential grooves. The grooves can be formed in the substrate in some embodiments.

“Process 2100 proceeds to act 2114 where a portion of a conductor will be wound within the first part of the grooves in the substrate. As we discussed in connection to FIG. 17A: In some embodiments, the first portion of conductor can be wound in alternating levels in grooves that are located at different levels, skipping each level. Other embodiments allow a portion (e.g. half turns) of a conductor to be wound within a portion (e.g. on each level) of the grooves while skipping the other parts. Act 2116 is then performed, in which a second portion would be wound within a second section of the grooves created in the substrate. The second part of the substrate, when wound from bottom to top of substrate, may be wound using parts of the grooves that were not used when the first portion of conductor was wound from top to bottom. The second section of the conductor can be wound from the bottom to its top using portions that cross over or under the first portion of the conductor. You can use any suitable conductor, including copper wire and litzwire. You can use single or multiple strands of conductor material. An amplifier may be attached to the ends of the conductor to amplify signals recorded by the RF coil when it is used in a low field MRI system to receive MR signals from imaged objects.

“FIGS. 22A-22L shows a process for manufacturing a transmit/receive RF coil that can be used in a low field MRI system. This is in accordance to some embodiments. FIG. FIG. 22A shows how the coil winding begins at the top of a substrate, e.g., a helmet made from plastic with grooves therein), in accordance to the numbered arrows. The conductor can be arranged in a connecting groove that connects the top of the substrate and a first circumferential slot. The conductor is then wound in a clockwise manner (2) around a half turn from the first circumferential groove. FIG. FIG. 22B shows that the conductor is placed (4) after the completion of (3) the half turn in the first circumferential slot. The connecting groove connects the first circumferential slot and a second groove which is further away from the top than it is in the first. The conductor is then wound around the opposite half turn of second circumferential groove in clockwise direction. FIG. FIG. 22C shows how the conductor is wound within the second circumferential slot until (7) a connecting groove is formed between the second and third circumferential slots. 22D. 22D. The conductor is then placed (8) in the connecting groove between second and third circumferential holes. FIG. FIG. 22E shows how winding (9) continues in third circumferential slot in clockwise direction, turning in opposite half-turn until (10) a connecting groove between the fourth and third circumferential slots is reached. As shown in FIG. 22F. 22F. The conductor is now placed (11) in the connecting groove between the fourth and third circumferential grooves. FIG. FIG. 22G shows how winding continues in the half-turn pattern described above up to the bottommost circumferential groove. As shown in FIG. 22, some embodiments do not allow the conductor to cross the helmet’s posterior side. 22H.”

“FIG. “FIG. 22I shows that, after winding the conductor in the bottommost circumferential slot is completed, winding continues from the bottom to the top within the areas of the circumferential slots that were not used during the winding from the top to the bottom. The conductor (12) is wound in the bottommost circular groove. The conductor (13) is then arranged within the connecting groove between the bottommost and upper circumferential grooves. The winding (14) continues in the circumferential groove that was left over from the top to the bottom. FIG. FIG. 22J shows that the winding (15), continues until the next connecting slot is found. The conductor is then arranged (16), and crosses over the conductor in the connecting groove. Finally, the conductor continues (17) in a higher circumferential groove. FIG. FIG. 22K shows how the winding continues to the top of substrate. After that, the conductor is removed to complete the coil winding for transmit/receive radio coil with interlaced winding as in FIG. 22L. 22L. The process shown in FIGS. FIGS. 22A-L depict winding half turns in each circumferential groove, but it is possible to use a different winding pattern.

“FIG. “FIG. 23A illustrates a method for making a receive-only radio frequency coil using an interlaced winding arrangement in accordance to some embodiments. The coil winding begins by placing (1) the conductor on top of the substrate (e.g. a plastic helmet with grooves) in a connecting groove on one side (e.g. the left side) of substrate. FIG. FIG. 23B shows how the conductor winds (3) around the groove when it reaches (2) the left-hand ring groove. FIG. FIG. 23C illustrates that after winding (4) the conductor around a ring groove, the conductor is arranged ((5)) and crosses over the conductor at the connecting groove between the ring slot and the top. FIG. 23D: Winding (6) continues in the left-hand half of the helmet in a curved groove. FIG. FIG. 23E illustrates that the conductor must be positioned to cross the top of a helmet after winding in the left helmet half is complete. This will allow the conductor to wind on the right helmet half as in FIG. 23F. FIGS. FIGS. 23G and 23H illustrate that the winding in right-hand helmets continues around the grooves of the right side of the helmet, and is arranged to cross over the conductor in connecting groove between the right-hand ring groove and the top.

“Having described various aspects and embodiments, it is evident that many modifications, modifications, or improvements to the technology disclosed in the disclosure will be possible to those who are skilled in the art. These modifications, alterations, and improvements are all within the scope and spirit of the technology described. People of ordinary skill in art can easily envision many other methods and/or structures to perform the function or obtain the results. Each variation and/or modification is considered to be within the scope and spirit of the embodiments. Many alternatives to the particular embodiments herein will be recognized by those skilled in the arts. It is understood, therefore, that the above-described embodiments are only examples and that inventive embodiments can be used within the scope and equivalents of the appended claims. Any combination of features, articles, material, kits, or methods described herein is also included in the scope of this disclosure, provided that such features, articles and materials, as well as any other features, systems and methods, are not mutually exclusive.

The above-described embodiments may be implemented in many ways. A variety of aspects and embodiments in the present disclosure that involve the performance of processes or methodologies may use program instructions executable on a device (e.g. a computer, processor, or another device) to perform or control the methods or processes. Many inventive concepts can be embodied in a computer readable medium (or multiple computer readable media), which encodes one or several programs that when executed on one of the computers or other processors perform one or many of the methods described above. Computer readable media or media can also be portable, so that programs stored on it can be loaded onto other computers or processors to implement different aspects. Computer readable media can be non-transitory in some instances.

“The terms ‘program? “Program?” or’software? are generic terms that can be used to refer to any type of computer code or set of instructions. “Software” is used in this context to mean any type of computer code, or set of instructions that can be used to program a computer to execute various aspects of the above. It should also be noted that one or more computer program that executes methods of this disclosure does not have to reside on one computer or processor. They may be distributed modularly among multiple computers or processors in order to implement different aspects of the disclosure.

Computer-executable instructions can be in many forms such as program modules. They are executed by one or more computers, or any other device. Program modules can include routines, programs and objects as well as data structures. Modules that are used to perform specific tasks or implement certain abstract data types. The functionality of program modules can be combined or distributed in different ways depending on the need.

“Data structures can also be stored on any computer-readable media. Data structures can be simplified by showing fields that can be related to their location within the data structure. These relationships can also be established by assigning storage to the fields that correspond with their locations on a computer-readable medium. Any mechanism that can establish a relationship between elements of a data structure may be used, such as pointers, tags, or any other mechanism that establishes relationship between them, could be used.

“The embodiments described above can be implemented in many ways. The embodiments can be implemented in hardware, software, or a combination of both. The software code can be executed on any processor or group of processors. This is true regardless of whether the processors are located in one computer or distributed across multiple computers. Any component or collection that performs the functions mentioned above can be considered a controller. You can implement a controller in many ways. For example, you could use dedicated hardware or general purpose hardware (e.g. one or more processors) to execute the functions. The controller may also be programmed with microcode or software.

“Moreover, it is important to understand that a computer can be embedded in any number of forms such as a rack-mounted, desktop, laptop, or tablet computer. These are just a few examples. A computer can also be embedded in devices that are not considered computers but have suitable processing capabilities. This includes a Personal Digital Assistant (PDA), smartphone, or other portable or fixed electronic device.

A computer can also have input and output devices. These devices can be used to provide a user interface, among other purposes. Printing machines or display screens can be used for visual output, as well as speakers or other sound-generating devices to present audio output. Keyboards and other input devices such as touchpads, digitizing tablets, and mice are all examples of devices that could be used to provide a user interface. Another example is speech recognition, which can be used to receive input information from a computer.

These computers can be connected to one or more networks, in any form, including a local network or a large area network such as an enterprise network and intelligent network (IN), or the Internet. These networks can be built on any technology and can operate according to any protocol. They may include wired networks, wireless networks, or fiber optic networks.

“Also as mentioned, certain aspects can be embodied in one or more methods. You can arrange the acts that make up the method in any way you like. It is possible to create embodiments in which the acts are performed in a different order than what is illustrated. This may allow for simultaneous performance of some acts, even though they are shown in sequential sequence in the illustrative embodiments.

“All definitions as defined and used herein should be understood to control dictionary definitions, definitions within documents incorporated by reference, and/or the ordinary meanings of defined terms.”

“The indefinite articles?a? “The indefinite articles?a?? If not clearly stated, the terms?an,’ and?an,? should be taken to refer to?at most one.

“The expression?and/or? “The phrase?and/or, as it is used in the specification and the claims, should mean?either/both? The elements so conjoined are elements that are conjunctively and disjunctively found in different cases. Multiple elements are listed with?and/or. Multiple elements listed with?and/or should be understood in the same way, i.e.?one? or?more? All elements are so connected. Optionally, other elements can be present than those specifically identified by the “and/or” clause. Regardless of whether they are related or not to the elements identified, other elements may optionally be present. As an example, the reference to “A” and/or “B” can be used with open-ended language like “comprising”. In one embodiment, A can be used to refer to only A (optionally including additional elements than B); in another embodiment to B only (optionally incorporating elements other than A); and in yet another embodiment to both A (optionally containing other elements); and so on.

“As used in the specification and the claims, the phrase “at least one” means at least one element. When referring to a list containing one or more elements, it should be understood that at least one element is selected from any of the elements in this list. However, not all elements are included in the list. This definition allows elements to optionally exist other than those elements that are specifically identified in the list of elements. Refers to elements that are related or not related to the element specifically identified. As an example, let’s say that?at least one? of the elements A or B is missing. Or, alternatively,?at minimum one of A and/or B? Or, equivalently,?at minimum one of A or B? can be used to refer to, in one embodiment to at minimum one, optionally containing more than one A (and optionally incorporating elements other than B); in another embodiment to at most one, optionally involving more than one B (and optionally incorporating elements other than A); in yet another embodiment to at least one (optionally including more that one A) and at least one (optionally including more B); etc.

“Also the terminology and phraseology used in this document are for description purposes only and should not be considered limiting. The use of the word?including? ?comprising,? Or?having? ?containing,? ?involving,? “Integrating” (and variations thereof) is intended to include the items listed below and their equivalents as well as any additional items.

“All transitional phrases, such as?comprising?, are included in the claims and the specification. ?including,? ?carrying,? ?having,? ?containing,? ?involving,? ?holding,? ?composed of,? The like and similar are to be understood as open-ended. This means that they can include but not limit. The transitional phrases ‘consisting of’ and?consisting essentially of are not allowed. Only the transitional phrases?consisting of? shall be either semi-closed or closed transitional phrases, depending on the context.

Summary for “Radio-frequency coil signal chains for a low field MRI system”

“Magnetic resonance imaging, (MRI), is an important imaging method for many applications. It is extensively used in clinical and research settings to create images of the insides of the human body. MRI relies on the detection of magnetic resonance signals (MR), which are electromagnetic waves emitted from atoms when they experience state changes due to applied electromagnetic fields. Nuclear magnetic resonance (NMR), for example, detects MR signals emitted by the nuclei excited atoms during the re-alignment of the nuclear spin of atoms in an object being scanned (e.g., atoms within the human tissue). The processing of MR signals can produce images. This is useful in medical applications. It allows the examination of internal structures and/or biological functions within the body for diagnostic, therapy, and/or research purposes.

“MRI is a popular imaging modality for biological imaging. It can produce images with high resolution and contrast that are non-invasive, without needing to expose subjects to ionizing radiation (e.g. x-rays or introducing radioactive material). MRI can also provide soft tissue contrast which can be used to image subjects that other imaging modalities cannot. MR techniques can also capture information about structures and/or biological process that other modalities cannot. There are a few drawbacks to MRI. These include the high cost of equipment, limited availability, difficulty in obtaining access to scanners for clinical MRI, and/or the lengthening of image acquisition.

The trend in clinical MRI is to increase the field strength to MRI scanners in order to improve scan time, image resolution and image contrast. This, in turn, drives up costs. Most MRI scanners are capable of operating at 1.5 to 3 tesla (T), which is the field strength of B0. An estimate of the cost for a clinical MRI scanner costs approximately one million dollars per tesla. This does not include the significant operation, maintenance, and service costs associated with such MRI scanners.

“Conventional high-field MRI systems require large superconducting magnetic magnets and associated electronics to create a strong uniform magnetic field (B0), in which an object (e.g. a patient) can be imaged. These systems are large, with a typical MRI installation having multiple rooms for the magnet and electronics, as well as the thermal management system and control console areas. MRI systems are expensive and large, so they can only be used by facilities such as hospitals or academic research centers that have the resources and space to buy and maintain them. MRI scanners are not readily available due to their high cost and large space requirements. There are many clinical situations where an MRI scan could be helpful. However, due to the above limitations, it is often not possible or impossible to do so.

“Some embodiments include a switch circuit that can be coupled to an RF coil of a low field magnetic resonance imaging system. The circuit includes at least one field effect transistor, (FET), that can be used as an RF switch at a frequency lower than 10 MHz.

“Some embodiments contain a drive circuit that applies a gate voltage at least to one field-effect transistor, (FET), which is used to operate as a radio frequency switch in a low magnetic resonance imaging system. The drive circuit includes at least one isolation element that isolates a voltage source from at least one FET.

“Some embodiments contain a circuit that tunes a radio frequency coil (RF) coupled to an amplifier of low-field magnetic resonance imaging systems. The circuit includes tuning circuitry that is coupled across inputs to the amplifier, as well as active feedback circuitry that is coupled between the output of an amplifier and an input.

“Some embodiments contain a circuit that tunes a radio frequency coil (RF) coupled to an amplifier of low-field magnetic resonance imaging systems. To reduce the quality factor of the RF coil, the circuit includes active feedback circuitry that is coupled between the output of an amplifier and an input to the amplifier.

“Some embodiments include a technique for tuning a radio frequency coil (RF) coupled to an amplifier of low-field magnetic resonance imaging systems. This involves arranging tuning circuitry at the first and second inputs of an amplifier and coupling active feedback between the output and input of an amplifier.

“Some embodiments include radio-frequency (RF), coils for low-field magnetic resonance imaging systems. The RF coil includes a substrate with a first and second sides, and a conductor that wraps around the substrate at a first plurality locations between the second and first sides. A second portion is wound around the substrate at a second plurality locations between the first and second sides.

“Some embodiments include a method for manufacturing a radio frequency (RF) coil that can be used in a low field magnetic resonance imaging system. This method involves providing a substrate with circumferential holes formed at multiple levels. Connecting grooves are placed at different distances from a first substrate side. Next, wind within a portion of circumferential and connecting grooves a first portion and second portions of conductors from the second substrate side to the first substrate side. The connecting grooves contain a second portion and a third portion of conductors.

“Some embodiments include an RF coil that can be used in low-field magnetic resonance imaging systems. The RF coil is composed of a substrate with a first and second side, and a conductor wrapped around the substrate in a balanced wounding pattern. In this pattern, the conductor wraps around the substrate from one side to another, crossing over the second section of conductor from the second side.

“The above-described apparatus and method embodiments can be used with any combination of features, aspects, and acts as described or further below. These, as well as other aspects, embodiments and features, can be better understood by reading the following description and referring to the accompanying drawings.

High-field systems dominate the MRI scanner market, especially for clinical or medical MRI applications. The general trend in medical imaging is to produce MRI scanners that have greater field strengths. Most clinical MRI scanners operate at 1.5T or 3T, with higher field strengths for research applications. As used herein, ?high-field? MRI systems currently in use in a clinical setting, and more specifically, MRI systems operating with a primary magnetic field (i.e. a B0 field) at 1.5 T or higher, although clinical systems operating between 0.5 and 1.5 T are also referred to as ‘high-field. Field strengths of between 0.2 T to 0.5 T are considered’mid-field. As field strengths in high-field have increased, so have field strengths between 0.2 T and 0.5 T. Field strengths between 0.5 T to 1 T have been described as mid-field. By contrast, ?low-field? Low-field MRI systems have a B0 value less than or equals approximately 0.2T. Systems with a B0 value between 0.2T and approximately 0.3T are sometimes referred to as low-field due to higher field strengths at high frequencies. Low-field MRI systems with a B0 field less than 0.1T are referred herein as “very low-field”. Low-field MRI systems with a B0 field less than 10 mT and lower-field systems are referred herein as ‘ultra-low field.

“Conventional MRI systems require specialized facilities, as we have already discussed. The MRI system must operate in an electromagnetically shielded area. Additionally, the room’s floor must be structurally reinforced. Additional rooms are required for high-power electronics as well as the control area of the scan technician. Access to the site must be secured. A dedicated three-phase electrical connection is required to supply power to the electronics. These electronics are then cooled using chilled water. Additional HVAC capacity is often required. These requirements not only make it costly but also limit the areas where MRI systems are possible to be installed. The operation and maintenance of conventional clinical MRI scanners requires a lot of expertise. The ongoing operational costs of operating an MRI scanner are high due to the need for highly-trained technicians and service engineers. Conventional MRI is expensive and difficult to access. This makes MRI a less accessible diagnostic tool that can deliver a broad range of clinical imaging solutions. MRI is not able to be used in many medical applications. It can assist with diagnosis, surgery and patient monitoring.

“High-field MRI systems, as we have discussed, require special adapted facilities that can accommodate their size, weight and power consumption. A 1.5 T MRI system weighs approximately 4-10 tons, while a 3T MRI system weighs about 8-20 tons. High-field MRI systems require a lot of expensive and heavy shielding. Mid-field scanners can weigh between 10-20 tonnes due to large permanent magnets or yokes. Low-field MRI systems, such as those that operate with a B0 magnetic fields of 0.2 T, are typically 10 to 20 tons in weight. This is due to the high amount of ferromagnetic materials used to generate the B0 field and additional shielding. Rooms must have a minimum of 30-50m2 to accommodate heavy equipment. They also need to have reinforced flooring (e.g. concrete flooring) and special shielding to protect the MRI system from electromagnetic radiation. Available clinical MRI systems cannot be moved and will require a dedicated space in a hospital or other facility. In addition to the high costs involved in preparing the space for operation and the ongoing costs associated with maintaining and operating the system, there will also be additional costs.

“In addition, MRI systems currently in use typically consume large amounts power. Common 1.5 T and 3T MRI systems consume 20-40 kW of power while 0.5 T and 0.2T MRI systems consume between 5-20 kW. Each system uses dedicated and specialized power sources. Power consumption refers to the average power used over a specified time period. The 20-40 kW referenced above is the average power consumed during image acquisition by conventional MRI systems. This may include short periods when peak power consumption is significantly greater than the average power consumption (e.g. when gradient coils or RF coils pulsed for short periods of time). The MRI system’s power storage elements (e.g. capacitors) can be used to address peak or large power consumption intervals. The average power consumption is therefore more important as it determines the type and frequency of the power connection required to operate the device. The available clinical MRI systems need to have dedicated power sources. This means that the MRI system will typically require a dedicated three phase connection to the grid. To convert the three-phase power into single phase power for the MRI system, additional electronics are required. Due to the many physical requirements for deploying conventional clinical MRI system, there is a limited availability of MRI and this severely limits the clinical applications that MRI can be used.

“Accordingly, high-field MRI installations are prohibitive in many cases. They limit their deployment to large institutions or specialized facilities. Also, they restrict their use to strictly scheduled appointments. Patients must visit designated facilities at times and places that have been scheduled in advance. High-field MRI is not fully used as an imaging modality due to the numerous restrictions. Despite these drawbacks, high-field MRI continues to be attractive due to the substantial increase in SNR at higher areas. This drives the industry to use higher field strengths in clinical and medical MRI applications. It also increases the cost and complexity for MRI scanners and limits their availability, preventing them from being used as a general-purpose or generally-available imaging solution.

“The low SNR of MR signal produced in low-field (especially in the very low field) has prevented the development and commercialization of a portable, cost-effective, low-power MRI system. Conventional ‘low-field? MRI systems are referred to as?low-field? because they operate at the highest end of what is commonly referred to as the low field range. (e.g. clinically available low fields have a floor of about 0.2 T) in order to produce useful images. Although they are less expensive than high-field MRI system, conventional low-field MRIs have many of the same problems. Low-field MRI systems, in particular, are heavy, immobile and expensive. They also consume significant power and need dedicated three-phase power hookups. Low-field MRI has made it difficult to develop portable, cost-effective, and/or low-power MRI systems that produce useful images.

“The inventors have created techniques that enable portable, low-field and low power MRI systems. This can increase the widespread deployment of MRI technology in a wide range of environments beyond those currently available at research and hospitals. MRI can now be used in emergency rooms, small clinics and doctor’s offices as well as in mobile units in the field. You can bring MRI equipment to your patient’s bedside to perform many imaging procedures. Some embodiments include low-field MRI systems (e.g. 0.1 T, 50 mT or 20 mT), These systems allow for portable, low-cost and low-power MRI. This greatly increases the availability of MRI in a clinical setting.

“Developing a clinical MRI system for low-field imaging poses many challenges. The term “clinical MRI system” is used herein to refer to an MRI that produces clinically-useful images. This means images with sufficient resolution and adequate acquisition time to be useful for a clinician or physician in order to fulfill a specific imaging application. As such, the resolutions/acquisition times of clinically useful images will depend on the purpose for which the images are being obtained. The relatively low SNR is one of the many challenges that can be faced when trying to obtain clinically useful images from low-field conditions. The relationship between SNR (and B0 field strength) is roughly B0 5/4 for field strengths above 0.2 T, and about B0 3/2 for field strengths below 0.01 T. This means that SNR decreases significantly with decreasing field strength, with even greater drops at low field strengths. The significant drop in SNR due to a decrease in field strength has been a major factor in the development of clinical MRI systems operating in the low-field range. The low SNR at low field strengths has made it difficult to develop a clinical MRI system that can operate in this low-field environment. Clinical MRI systems that aim to operate in lower field strengths have traditionally achieved field strengths of around 0.2 T and higher. These MRI systems can be heavy, expensive, and large. They require dedicated space (or shielded tents), and power sources.

“Inventors have created low-field and very-low-field MRI systems that can produce clinically useful images. This allows for the creation of portable, affordable, and easy-to-use MRI systems not possible with state-of-the-art technology. An MRI system can be carried to the patient in some embodiments to offer a variety of diagnostic, surgical and monitoring procedures.

“FIG. 1. A block diagram showing typical components of a MRI 100 system. FIG. 1. MRI system 100 consists of computing device 104 and controller 106. It also includes pulse sequences store 108, power management 110, magnetics components 120, and power management system 110. System 100 is an illustration. A MRI system could have additional components of any type, in addition or instead of those shown in FIG. 1. These high-level components will be included in an MRI system, but the implementation of those components may vary for each MRI system. We’ll discuss this further below.

“As illustrated at FIG. 1. Magnetics components 120 include B0 magnets 122, shims 124 and RF transmit-receive coils 126, and gradient coils 128. Magnet 122 can be used to create the main magnetic field B0. Magnet 122 can be any combination or type of magnetic components that generates the desired main magnetic field B0. The B0 magnet, which is formed in high-field conditions, is usually made using superconducting material, generally in a solenoid geometrie. This requires cryogenic cooling systems to maintain the superconducting state of the magnet. High-field B0 magnets can be expensive and complicated, consume large amounts of energy, and are therefore costly and complex. Superconducting material is also used to implement conventional low-field B0 magnetics (e.g. B0 magnets operating at 0.2% T). To produce the low field strengths of conventional low-field B0 magnetics (e.g. at 0.2 T), permanent magnets must be used. These magnets need to weigh 5-20 tons to achieve the required field strengths. The B0 magnet in conventional MRI systems is not portable and affordable.

Gradient coils 128 can be arranged in a way that generates gradient fields. For example, they may be arranged in three orthogonal directions (XYYZ) in order to create gradients in B0’s field. Gradient coils 128 can be used to encode emitted MR signal by systematically varying B0 fields (the B0 field generated from magnet 122 or shim coils 124) in order to encode the spatial position of received MR signals as a function frequency or phase. Gradient coils 128 can be set up to change frequency or phase in a linear function. However, nonlinear gradient coils may allow for more complex spatial encoding profiles. A first gradient coil might be designed to selectively alter the B0 fields in a first direction (X), to perform frequency encoding in that direction. A second gradient coil could be set up to selectively change the B0 fields in a second direction (Y), substantially orthogonal, to perform phase encoders. A third gradient coil could be set up to selectively adjust the B0 fields in a third direction (Z), substantially orthogonal, to allow slice selection for volumetric imaging applications. The conventional gradient coils consume considerable power and are typically powered by expensive, large-scale gradient power sources. We will discuss this further below.

“MRI is achieved by stimulating and detecting emitted MR signal using transmit and/or receive coils (often referred as radio frequency (RF), coils). Transmit/receive coils can include separate coils to transmit and receive, multiple coils to transmit and/or receive, or the same coils to transmit and receive. A transmit/receive element may contain one or more coils to transmit, one or several coils to receive, and/or one/more coils to transmitting or receiving. To refer to the different configurations of the transmit/receive magnetics components of an MRI system, transmit/receive coils may also be called Tx/Rx. These terms can be interchanged herein. FIG. FIG. 1. RF transmit and receiver coils 126 are made up of one or more transmit coils. These coils can be used to generate RF pulses that induce an oscillating magnet field B1. You can configure the transmit coils to produce any type of RF pulses.

“Power management system 110” includes electronics that provide power for one or more low-field MRI systems 100 components. As an example, power management system 110 could include power supplies, transmit coil components and/or other power electronics to provide sufficient operating power to energize or operate components of MRI 100. FIG. FIG. The electronics in power supply 112 provide the operating power for magnetic components 120 of MRI system 100. Power supply 112 could include electronics that provide power to B0 coils (e.g. B0 magnet 122) in order to generate the main magnetic field for low-field MRI systems. The transmit/receive switch (116) can be used to control whether RF transmit or RF receive coils will be operated.

“Power component(s), 114 may contain one or several RF receiver (Rx) preamplifiers that amplify MR signal detected by one/more RF transmitter coils. (e.g. coils 126), one (or more) RF transmit power components designed to supply power to one (or more) RF transmit coils. (e.g. coils 126), one (or more) gradient power components which provide power for one (e.g. gradient coils 128) and one (e.

“Power components in conventional MRI systems are expensive, large and use significant power. The power electronics are usually located in a separate room from the MRI scanner. Power electronics require a lot of space. They are complex and expensive devices that require support from wall-mounted racks. The power electronics in conventional MRI systems prevent portability and affordability of MRI.

“As shown in FIG. “As illustrated in FIG. 1, MRI system 100 comprises controller 106, also known as a console. It has control electronics that send and receive instructions from power management system 110. The controller 106 can be configured to execute one or more pulse sequences. These are used to determine instructions to power management system 110 to control magnetic components 120 according to a desired sequence (e.g. parameters to operate RF transmit and receiver coils 126, operating gradient coils 128, etc.). FIG. FIG. Computing device 104, for example, may use received MR data to create one or more MR images by using any image reconstruction process. For data processing by the computing device, controller 106 might provide information about one or several pulse sequences to computing devices 104. Controller 106, for example, may give information to computing device104 about one or more pulse sequences. The computing device may then perform an image reconstruction process, at least partially, based on this information. Computing device 104 is typically a high-performance workstation that can perform complex computationally intensive processing on MR data quickly. These computing devices can be quite expensive.

“As you can see, the current clinical MRI systems (including mid-field, high-field, and low-field) are expensive, large, fixed systems that require large dedicated spaces and dedicated power connections. The inventors developed low-field, and very-low-field MRI systems, which are more cost-effective, less powerful, and/or portable. This greatly increases the accessibility and application of MRI. Some embodiments provide a portable MRI system that can be carried to patients and used at the locations it is most needed.

“Some embodiments of MRI systems are portable. This allows the MRI device, as well as the corresponding power consumption, to be transported to the locations where it is needed. The development of a portable MRI device is not without its challenges. It must be small and light, consume little power, and operate in uncontrolled electromagnetic noise environments (e.g. outside a special shielded area).

“Portability refers to the ability to operate the MRI system in a variety of environments and locations. The current clinical MRI scanners must be installed in specially protected rooms. This is, among other reasons, why they are not portable, non-availability, and cost prohibitive. The MRI system must be able to operate in a wide range of noise environments, and therefore, cannot be used outside of a specially protected room. The inventors have developed noise suppression techniques that allow the MRI system to be operated outside of specially shielded rooms, facilitating both portable/transportable MRI as well as fixed MRI installments that do not require specially shielded rooms. These noise suppression techniques are able to be used outside of specially shielded rooms. However, they can also be used in shielded environments.

The MRI system’s power consumption is another aspect of portability. As mentioned above, clinical MRI systems use a lot of power. They consume between 20 kW and 40 kW per hour. This means that dedicated power connections are required. These connections can be three-phase power connections or dedicated connections to the grid. A dedicated power connection is required to operate an MRI system in other locations than costly dedicated rooms. Low power MRI systems have been developed by the inventors. They can be used with mains electricity, such as a 120V/20A connection in the U.S. or large appliance outlets (e.g. 220-240V/30A), allowing the device’s operation wherever common power outlets are available. You can plug into the wall. facilitates both portable/transportable MRI as well as fixed MRI system installations without requiring special, dedicated power such as a three-phase power connection.”

“A portable MRI device is designed according to the techniques discussed herein and includes RF transmit coils 126 that generate a B1 magnetic fields during a transmit operation and collect flux from an MR signal created by an imaged object during a receieve operation. The RF receive coil amplifies and processes signals before they are converted into MR images. The RF signal chain is the circuitry that controls and processes signals recorded by the coils 126. circuitry. The inventors recognized that the components of the RF signals chain circuitry used in high-field MRI systems are not suitable and/or optimized to be used in low-field MRI systems designed in accordance herewith. Some embodiments provide improved RF signal chains circuitry that can be used in a portable, low-field MRI system.

“FIG. “FIG. The RF signal circuitry 200 comprises RF transmit/receive 210 and transmit/receive 212. These circuitry are used to selectively couple the RF receiver circuitry to the RF coils 210 depending on whether they are being used to transmit or receive. The Larmor frequency is a frequency that RF coils should resonate at in order to perform optimally. According to the following relation, the Larmor frequency (w), is related with the strength of B0 field. The gyromagnetic relationship of the imaged element (e.g. 1H) in MHz/T is the gyromagnetic value of the isotope, while B is the strength the Tesla’s B0 field is in Tesla. For a 1.5T MRI system, the Larmor frequency is approximately 64 MHz and for a 3T MRI system it’s approximately 128 MHz. The Larmor frequency for low-field MRI is significantly lower than that of high-field MRI. The Larmor frequency of a 64 mT MRI is 2.75 MHz. RF signal chain circuitry 200 also includes tuning/matching (214) circuitry that transforms the impedance RF coil 210 to maximize performance. Amplifier 216, which is a low-noise amplifier, receives the output of tuning/matching 214. It amplifies RF signals before they are converted into image signals. One of the problems with low-field MRI systems using RF coils is their susceptibility to noise in electronic parts. According to some embodiments, components 210, 214, and 216 can be configured to reduce noise in RF signal chains.

“Some embodiments contain multiple RF coils in order to increase the signal-to noise ratio (SNR), of signals detected by an RFID coil network. A collection of RF coils can be placed at different orientations and locations to detect an extensive RF field. To improve image acquisition’s SNR, a portable MRI system may include multiple RF transmit/receive coils. A portable MRI system could include 2, 4, 8, 16, 32, or more RF receiver coils to increase the SNR for MR signal detection.

“RF coils can be tuned to increase the frequency at which they are sensitive (e.g. the Larmor frequency). Inductive coupling between neighboring or adjacent coils (e.g., coils located sufficiently close to one another) reduces the sensitivity and dramatically reduces the effectiveness for the collection of RF coils. There are techniques for geometrically decoupling neighboring coils, but these places strict restrictions on the coil position and orientation in space. This reduces the ability to detect the RF field accurately and decreases the signal-to noise performance.

“To reduce the negative effects of inductive co-upling between coils, inventors used coil decoupling techniques to reduce the inductive coupling effect between radio frequency coils in multicoil transmit/receive system. FIG. 3 illustrates an example of a passive decoupling circuit 300 that reduces inductive coupling between radio frequency coils in multi-coil transmit/receive systems. FIG. 3 shows a passive coupling circuit 300 that reduces inductive coupling between radio frequency coils in multi-coil transmit/receive systems. Circuit 300 can be used to decouple RF coils exposed to B1 transmit frequencies (e.g. from an RF transmit loop). The decoupling circuit reduces the current flowing through an RF coil when there is an AC excitation voltage at the Larmor frequency. Inductor L1 is an RF signal coil that is visible to the MRI system’s field of view. To optimize noise performance impedance, capacitors C1 and C2 create a tuning circuit which matches the inductance to the input of low noise amplifier (LNA). To prevent the RF coil from being paired with other coils, the tank circuit which includes capacitor C3 and inductor L2 is reduced by the capacitor C3 and capacitor L1. FIG. FIG. 4A shows a plot of voltage at LNA input at resonant frequency for RF coil. This plot is based on simulation of circuit 300 in FIG. 3. FIG. 4B shows a plot of current through an RF coil, based on simulations of circuit 300 in FIG. 3. The LNA voltage at 2.75 MHz resonant frequency is approximately 26 dB (FIG. 4A and the coil current are?37 dB (FIG. 4B). FIGS. FIGS. 4A and 4B show the magnitude of the measured quantity as a straight line, while the phase of that quantity is shown as a dashed or dashed line.

The inventors recognized that decoupling with a tuned matching filter to lower the current in an RF coil has its drawbacks. For example, it requires tuning multiple components (e.g. capacitors C1,C2 and C3) to the operating frequency. SNR is also affected by losses in L2 inductor. Decoupling efficiency, therefore, is a compromise between SNR efficiency and decoupling efficiency. FIG. FIG. 4B shows that although the tuned match filter reduces coil current substantially at the resonance frequency, the sharp valley of the current waveform indicates that the current reduction through RF coil is very limited for the narrow bandwidth around the resonant frequencies.”

“Some embodiments refer to an improved decoupling system that reduces the current in the RF coil. This is done by dampening the coil response with feedback from the amplifier. FIG. FIG. 5 shows a circuit 400 that provides feedback decoupling according to some embodiments. Circuit 400 has an active feedback path that runs from an amplifier LNA’s output to an input LNA. FIG. 5 shows an example of this active feedback path. FIG. 5 shows the active feedback path. It includes one feedback path. It should be noted that the active feedback path can be implemented in a variety of ways, each providing a different type or feedback decoupling depending on which is selected. In some embodiments, for example, the active feedback path may include a first and second feedback paths that provide feedback signals.

The phase of the feedback signal has an effect on the tuning frequency’s amplification gain. This was recognized by the inventors. In some embodiments, multiple feedback paths are used in the active feedback path. A first feedback path might provide a 90-270 degree out of phase feedback signal with a frequency of the coil, while a second feedback pathway may provide a 180-degree out of phase feedback signal with a frequency of coil. Alternately, the gain of an amplifier can be tuned to be 90 degrees or 270 degrees out-of-phase with the resonant frequency. The maximum amplification gain at the tuning frequencies may be achieved when the phase is 270 degrees. Other embodiments may use a single feedback path. In these cases, the phase of feedback signal can be adjusted to 180 degrees to achieve more efficient decoupling.

Circuit 400 provides feedback decoupling by using active negative feedback to dampen the coil response. Also known as reducing Q factor (or?deQing?). The coil) and reduces the current flowing in it. Circuit 400 includes a tuning/matching loop between the LNA and the RF coil. You can use any suitable tuning/matching device in accordance to certain embodiments. Examples are given below.

“FIG. “FIG. Decoupling circuit 500 uses only one component, i.e. capacitor C1, to tune, in contrast to decoupling Circuit 300. Circuit 500 does not contain reactive components C1 or C2 and therefore does not have an inductor. This prevents the SNR losses that are associated with circuit 300 (inductor L1 being in circuit 300).

“Capacitor C1 can be implemented with a capacitor of fixed value. Alternately, capacitor C1 can be implemented with a capacitor having a fixed value (e.g., via a varactor diode). Another way to implement capacitor C1 is to use a capacitor of fixed value (e.g. 300 pF) in parallel with a capacitor of variable value. This arrangement reduces AC losses caused by the use of a variable capacitor within the feedback loop.

“FIG. 7A shows a plot of voltage at LNA input at resonant frequency RF coil. This is based on simulations of circuit 500 in FIG. 5. FIG. FIG. 7B shows a plot of current through RF coil, based on simulation of circuit 500 in FIG. 6. The LNA input voltage at 2.75 MHz resonant frequency is approximately 8 dB. (FIG. 7B, and the current through the coil at?35 dB. Contrary to FIG. FIG. 7B shows that coil current is lower when decoupling circuit 500 is used than circuit 300. Circuit 500, by contrast to circuit 300, provides RF coil-decoupling over a larger bandwidth.

“FIG. “FIG.8” illustrates an alternate feedback-based decoupling circuit 600. In this circuit, the single capacitor tuning/matching loop of circuit 500 is replaced by an alternative feedback circuit. 6 is replaced by a tuning/matching system that includes components C1,C3 and L2. Circuit 600 uses a tuning/matching system to tune the RF coil (represented by L1), in addition to feedback-based decoupling via an active feedback path that includes capacitor C2.

“In certain embodiments, capacitive feedback circuitry is provided, for instance, by the feedback parts of circuits 400 to 500 and 600 in FIGS. 5 and 8 are replaced by mutual inductive feedback circuitry. FIG. FIG. 24 shows an alternative feedback-based circuit 2400. In this circuit, the capacitive feedback circuitry of circuit 500 is replaced by one that uses feedback. 6 is replaced by mutual inductive feedback circuitry, which includes components R1, L2 and R2. Circuit 2400 has inductors L1 & L2 that are connected mutually using a transformer, or by the air.

A transmit/receive switch is another technique to achieve RF coil decoupling according to some embodiments. When RF signals are being transmitted from one or more RF transmit loops, the transmit/receive toggle switch isolates the RF coil and the amplifier. The transmit/receive toggle divides the tuning/matching networks into two sections to protect sensitive electronics during RF transmit cycle. The transmit/receive button 312 in some conventional MRI systems, such as high-field MRI system, is usually implemented with a diode like a PIN diode. FIG. 9 shows a diagram of transmit/receive circuitry with a diode, D1. 9 is circuit 700. The transmit pulse causes diode D1 to be turned on. This creates a short circuit that isolates the RF signal coil and the receive electronics. The resulting network, as described in connection to circuit 300, creates a tank circuit that has a high impedance. This ensures that the current in RF coil stays small. The RF coil is connected to the amplifier during receive cycles. This allows the tank circuit to tune the RF coil to reduce the current flowing through it while still allowing sufficient signal to reach the amplifier’s output. The RF coil is connected during transmit cycles to a first tank circuit and a second tank loop during receive cycles.

“Conventional decoupling systems, such as the one shown in FIG. PIN diodes are often used to isolate the receiver electronics from the RF signal loop in conventional decoupling circuits, such as the one shown in FIG. To turn on a PIN diode in a decoupling circuit, however, it requires approximately 0.1 A. For example, a transmit/receive system with eight coils might require 0.8 A to decouple each receive coil from the RF signal coil for each transmit or receive cycle in an image acquisition pulse sequence. The decoupling circuits in the RF transmit/receive systems consume significant power over the course of an image acquisition protocol. When PIN diodes have been used, a biasing resistor (R1) and an AC blocking filter consisting of components L1 & C1 are required. The ground circuit is not isolated when it is in its off state. While PIN diodes are able to work at higher frequencies in high-field MRI systems (e.g. 10 MHz), they do not perform well at lower operating frequencies, such as those used in very low field or low-field MRI systems. The PIN diode works by rectifying the signal, not blocking it at such low frequencies. A DC bias current Ibias, for example, allows the diode’s forward bias to continue even when a negative signal has been applied. To block an AC signal with frequency f and peak current Ipeak the ratio Ipeak/f must be lower than the sum of the DC bias current Ibias, and the carrier life t. Some low-field MRI applications might have these parameters: Ipeak=10 A. f?2.75 MHz. Ibias=100 mA. According to the relation above, for these parameters, the PIN diode would need to have a carrier lifetime ?>37 ?s, which is not a characteristics of commercially-available PIN diodes.”

“Inventors recognized that PIN Diodes commonly used in a decoupling system may be replaced with Gallium Nitride GaN field effect transistors (FETs). This addresses some of the limitations of using PIN DIodes in RF transmit/receive systems of low-field MRI systems, including reducing power consumption. GaN FETs are much more efficient than PIN diodes, as they require only a few microamps to switch on. This reduces the power consumption by many orders of magnitude. The resistance of the GaN FETs is smaller than PIN diodes when they are turned on, which reduces the tank circuit’s impact. In some embodiments, diode 700 D1 is replaced by one or more GaN FETs to reduce the power consumption of RF transmit/receive systems.

“FIG. “FIG. Circuit 412 is composed of two mirrored FETs. However, some embodiments allow for an RF transmit/receive circuit 412 to include any number of FETs. GaN FETs are more efficient than PIN diodes. They operate at all frequencies and consume very little power.

“FIGS. 11A-C show operating states of FETs used as switches in RF transmit/receive systems in accordance to some embodiments. FIG. FIG. 11A shows a GaN FET that is configured as a switch between two drain nodes D and source nodes S. The gate of the GaN FET controls the state of the switch from on to off. FIG. FIG. 11.B shows that the GaN FET can also be modelled with three lumped capacitors: C_dss, C_gss and C_gd in the off state. The drain D can be isolated from the source S in such a configuration provided that C_ds is low (e.g. 10-100 pF). Some embodiments have a drain-source capacitance greater than 15 pF. FIG. FIG. 11.C shows that the drain-source capacitance (C_ds) is replaced in the on state by a short circuit.

“FIG. “FIG. 12” illustrates a circuit 1000 that drives a gate voltage on GaN transistors U1 or U2 and is arranged to act as an RF transmit/receive switching in accordance with certain embodiments. GaN FETs can be used to decouple and couple the receive electronics to the RF coil. Inductors L5 & L6 can be used as transformers to connect a control signal V2 with the gates of the FETs U1 & U2, and provide ground isolation. The diode (D3) rectifies the control signal to create DC on/off voltage across the capacitor C7. The resistor R11 can be used to discharge capacitor C7 as well as the gate capacitance for the FETs. The transmit/receive switch will turn off or on quickly depending on the time constants of C7+Cgates and R11. In certain embodiments, V2 can be a 10MHz sine wave that is coupled to L5 to drive FETs. The 10 MHz signal can be used to turn on/off the FET gates, and then switched off. To open the switch, R11 is used to discharge the gate drive resistor. FIG. 12 illustrates how the coupling between L5/L6 and inductors L5/L6 can be poor. Inductances may also be small. In some cases, L5/L6 could be used as a small air-core transformer (or as an RF transformer).

“FIG. “FIG. 13” illustrates a circuit 1100 that drives a gate voltage on GaN transistors U1 or U2. It is arranged to function as an RF transmit/receive switching in accordance with certain embodiments. Circuit 1100 uses the RF transmit pulse as the control signal to activate the transmit/receive switch, rather than an externally supplied control signal V2 like circuit 1000. FIG. FIG. 13 shows a coil that is represented by an inductor L6. This coil is designed to receive the RF transmitter pulse and generate a voltage which drives the GaN FETs. One embodiment may associate each of the RF coils with a coil L6 that receives the RF transmit pulse. Another embodiment may associate a subset of RF coils with a coil L6 that is configured to receive an RF transmit pulse. The switch signal generated in response by coil L6 may then be distributed to other circuitry in the array. Circuit 1000 does not require a separate control signal generator, so circuit 1100 has fewer complex circuitry. The transmit pulse is used as a control signal. However, the switch doesn’t close until RF transmission starts. Pulse receiver coil L6 can detect the RF transmit pulse.

“Some embodiments relate to a novel design of a radio frequency (RF) coil that can be used in a low field MRI system. Conventional RF coil designs are designed to work in MRI systems as a solenoid. This wraps around the object and creates a helix pattern. Head coils are used in MRI systems because they have a conductor that is in a solenoid arrangement. This allows for a head to be inserted into the solenoid. FIG. FIG. 14A shows a schematic illustration of a solenoid-RF coil design, in which a conductor wraps in a series of loops around the substrate in one pass. The loops are from the substrate’s first side to the second side. The conductor can be returned to its original side when the substrate’s second side is reached. FIG. FIG. 14B shows a top-view of the coil arrangement in FIG. 14A shows the conductor loops as vertical lines. Points V+ & V? V+ and V? are the ends of the conductors in the coil. They are connected to an amplifier in an MRI system (e.g. a low-noise amplifier) to amplify the recorded signals.

“In ideal cases, the potentials recorded at the outputs the RF coil are balanced so that V+/V?=0 in absence of electromotive force(emf). If an object such as the head or body of a person is inserted into the solenoid, parasitic coupling may occur between the object and conductor in the coil. This could lead to V+ and V? Unbalanced, resulting in a voltage at amplifier input. When the coil is used in an MRI system, the voltage is expressed as a noise signal in MR signal. The parasitic coupling can affect signals at V+ and V depending on where the head is located within the RF coil. differently. If the object is placed at one end, the amount of noise introduced by parasitic coupling into the recorded signal may be greater at V+ than at V. Because of the shorter conductor distance that separates V+ from the point at which noise was introduced into the coil, Alternately, if the object is placed at or near the centre of the coil, between V+, V?, then the noise introduced into the coil will affect both V+, V?. equally. Another implementation would see more noise at V if the object were placed closer to V?. ”

“FIG. 15A shows schematically that an object (represented as voltage source) is placed into a solenoid coil in a specific location. A parasitic coupling (represented as impedanceZC) is then introduced into the coil at one point. In practice, the parasitic coupling between the object and the coil winding will not be distributed. FIG. FIG. 15B shows an impedance model showing how parasitic coupling can affect the voltages V+ and V? The conductor’s ends. ZC is the parasitic coupling of the object and coil, Z+ the impedance within the conductor between where the parasitic coupling was introduced and Z+, Z? ZC represents the impedance of the conductor between V+, Z? and the point where the parasitic coupling was introduced. ZG represents each end of the conductor (i.e. V+ and V?). Ground. If there is weak parasitic coupling between an object and the coil (e.g. Zc?. Z+. Z?, ZG), then the following relation describes the difference of potential at the two ends V+ and V? :”

“V + V – =Z G Z C 2? ( Z ? – Z+ ) V 0

“Because V+ and V are the outputs of the coil? Some RF coils have a balun between their RF coil and amplifier in order to balance the output and reject common mode noise. Baluns are not recommended for low-field MRI systems because of the small magnitude signals received by the coil as well as the lossy characteristics that baluns can have. Some embodiments address an RF coil design which uses a winding pattern to reduce common mode noise. This eliminates the need for a balun.

“FIGS. 16-19 illustrate schematically RF coil designs according to some embodiments. FIGS. The result is a solenoid coil with similar magnetic properties to the coil designs shown in FIGS. 14A. 14A. For example, the magnetic flux detected by turns of the coil placed close together is similar. The electrical properties of the RF coil designs in FIGS. 16-19 are different. The electrical properties of the RF coil designs shown in FIGS. 16-19 are not the same as those shown in FIG. 14A coil design. Particularly, FIGS. 16-19 show the winding designs. FIGS. 16-19 show winding designs that improve balance and reject common mode when images are inserted into the coil. The parasitic coupling between an object and the conductor causes a voltage to be induced near the object when the object is inserted into a coil. FIGS. 16-19 show the winding patterns. FIGS. 16-19 show that adjacent turns of the conductor have the same inductance/potential when a voltage to them is applied. This is because they are at a similar distance from V+ and V??. The voltage caused by parasitic coupling an object in the coil to the conductor creates a noise signal similar to V+ and V?, regardless of its location in the coil. Common mode noise is reduced by?0

Instead of winding the conductor in a single loop from one end to the other, as shown in FIG. 14A: In some embodiments, the conductor can be wound in a balanced manner using multiple passes of loops (e.g. two or more) from one end to the other. FIG. FIG. 16A illustrates an “interlaced” arrangement. 16A shows an “interlaced” winding pattern. A conductor is wound around a substrate starting at one end. The conductor then moves along the winding direction in a single pass, skipping different portions of the substrate at different levels. The conductor is wound around the substrate portions that were missed in the first pass when it is wound from the second (other) end of the substrate. FIG. FIG. 16B is a top view showing the interlaced winding pattern shown in FIG. 16A.”

“FIG. 18A shows an alternate balanced winding pattern that conforms to some embodiments. FIG. FIG. 18A shows a first plurality loops of conductor that are wound around the substrate, passing from the first to the second ends without skipping any levels. To create a “double?”, a second plurality loops near the first plurality are wound around substrate on the second pass. Winding pattern. FIG. FIG. 18B shows a top-view of the winding pattern shown in FIG. 18A.”

“FIG. 19 is a top view showing another balanced winding configuration with an interlaced configuration, in accordance to some embodiments. FIG. FIG. 19 shows a winding pattern that forms loops at a series of levels, rather than from one end of a substrate to the other (e.g. as shown in FIG. 16A) The conductor is wound in two passes from the first to the second ends of the substrate. A first pass forms loops at a series of levels from one end to another end of the substrate (e.g., as shown in FIG. The invention does not limit the use of a particular angle for winding the helix. Any angle that has a desired number of turns around a substrate can be used.

“FIG. 21 shows a process 2100 to manufacture an RF coil according to some embodiments. Act 2110 provides a substrate around which the conductor is to wind. Any suitable non-magnetic material may be used as the substrate. The substrate may be made of a plastic material that has been fabricated using additive manufacturing processes (e.g. 3D printing). The process 2130 proceeds to act 2112 where a plurality grooves are created in the substrate. The substrate may have a top and bottom, and the plurality grooves can be placed at different locations from the top to the bottom. The substrate may be shaped like a helmet, where the head of an individual can be placed. In some embodiments, the grooves are made as a plurality circumferential grooves. From the top to bottom, the circumference of helmet. Some embodiments of the plurality rings are separated by the same spacing from top to bottom of substrate to create multiple levels in which a conductor can be wound. A plurality connecting grooves can also be included between the circumferential grooves. The grooves can be formed in the substrate in some embodiments.

“Process 2100 proceeds to act 2114 where a portion of a conductor will be wound within the first part of the grooves in the substrate. As we discussed in connection to FIG. 17A: In some embodiments, the first portion of conductor can be wound in alternating levels in grooves that are located at different levels, skipping each level. Other embodiments allow a portion (e.g. half turns) of a conductor to be wound within a portion (e.g. on each level) of the grooves while skipping the other parts. Act 2116 is then performed, in which a second portion would be wound within a second section of the grooves created in the substrate. The second part of the substrate, when wound from bottom to top of substrate, may be wound using parts of the grooves that were not used when the first portion of conductor was wound from top to bottom. The second section of the conductor can be wound from the bottom to its top using portions that cross over or under the first portion of the conductor. You can use any suitable conductor, including copper wire and litzwire. You can use single or multiple strands of conductor material. An amplifier may be attached to the ends of the conductor to amplify signals recorded by the RF coil when it is used in a low field MRI system to receive MR signals from imaged objects.

“FIGS. 22A-22L shows a process for manufacturing a transmit/receive RF coil that can be used in a low field MRI system. This is in accordance to some embodiments. FIG. FIG. 22A shows how the coil winding begins at the top of a substrate, e.g., a helmet made from plastic with grooves therein), in accordance to the numbered arrows. The conductor can be arranged in a connecting groove that connects the top of the substrate and a first circumferential slot. The conductor is then wound in a clockwise manner (2) around a half turn from the first circumferential groove. FIG. FIG. 22B shows that the conductor is placed (4) after the completion of (3) the half turn in the first circumferential slot. The connecting groove connects the first circumferential slot and a second groove which is further away from the top than it is in the first. The conductor is then wound around the opposite half turn of second circumferential groove in clockwise direction. FIG. FIG. 22C shows how the conductor is wound within the second circumferential slot until (7) a connecting groove is formed between the second and third circumferential slots. 22D. 22D. The conductor is then placed (8) in the connecting groove between second and third circumferential holes. FIG. FIG. 22E shows how winding (9) continues in third circumferential slot in clockwise direction, turning in opposite half-turn until (10) a connecting groove between the fourth and third circumferential slots is reached. As shown in FIG. 22F. 22F. The conductor is now placed (11) in the connecting groove between the fourth and third circumferential grooves. FIG. FIG. 22G shows how winding continues in the half-turn pattern described above up to the bottommost circumferential groove. As shown in FIG. 22, some embodiments do not allow the conductor to cross the helmet’s posterior side. 22H.”

“FIG. “FIG. 22I shows that, after winding the conductor in the bottommost circumferential slot is completed, winding continues from the bottom to the top within the areas of the circumferential slots that were not used during the winding from the top to the bottom. The conductor (12) is wound in the bottommost circular groove. The conductor (13) is then arranged within the connecting groove between the bottommost and upper circumferential grooves. The winding (14) continues in the circumferential groove that was left over from the top to the bottom. FIG. FIG. 22J shows that the winding (15), continues until the next connecting slot is found. The conductor is then arranged (16), and crosses over the conductor in the connecting groove. Finally, the conductor continues (17) in a higher circumferential groove. FIG. FIG. 22K shows how the winding continues to the top of substrate. After that, the conductor is removed to complete the coil winding for transmit/receive radio coil with interlaced winding as in FIG. 22L. 22L. The process shown in FIGS. FIGS. 22A-L depict winding half turns in each circumferential groove, but it is possible to use a different winding pattern.

“FIG. “FIG. 23A illustrates a method for making a receive-only radio frequency coil using an interlaced winding arrangement in accordance to some embodiments. The coil winding begins by placing (1) the conductor on top of the substrate (e.g. a plastic helmet with grooves) in a connecting groove on one side (e.g. the left side) of substrate. FIG. FIG. 23B shows how the conductor winds (3) around the groove when it reaches (2) the left-hand ring groove. FIG. FIG. 23C illustrates that after winding (4) the conductor around a ring groove, the conductor is arranged ((5)) and crosses over the conductor at the connecting groove between the ring slot and the top. FIG. 23D: Winding (6) continues in the left-hand half of the helmet in a curved groove. FIG. FIG. 23E illustrates that the conductor must be positioned to cross the top of a helmet after winding in the left helmet half is complete. This will allow the conductor to wind on the right helmet half as in FIG. 23F. FIGS. FIGS. 23G and 23H illustrate that the winding in right-hand helmets continues around the grooves of the right side of the helmet, and is arranged to cross over the conductor in connecting groove between the right-hand ring groove and the top.

“Having described various aspects and embodiments, it is evident that many modifications, modifications, or improvements to the technology disclosed in the disclosure will be possible to those who are skilled in the art. These modifications, alterations, and improvements are all within the scope and spirit of the technology described. People of ordinary skill in art can easily envision many other methods and/or structures to perform the function or obtain the results. Each variation and/or modification is considered to be within the scope and spirit of the embodiments. Many alternatives to the particular embodiments herein will be recognized by those skilled in the arts. It is understood, therefore, that the above-described embodiments are only examples and that inventive embodiments can be used within the scope and equivalents of the appended claims. Any combination of features, articles, material, kits, or methods described herein is also included in the scope of this disclosure, provided that such features, articles and materials, as well as any other features, systems and methods, are not mutually exclusive.

The above-described embodiments may be implemented in many ways. A variety of aspects and embodiments in the present disclosure that involve the performance of processes or methodologies may use program instructions executable on a device (e.g. a computer, processor, or another device) to perform or control the methods or processes. Many inventive concepts can be embodied in a computer readable medium (or multiple computer readable media), which encodes one or several programs that when executed on one of the computers or other processors perform one or many of the methods described above. Computer readable media or media can also be portable, so that programs stored on it can be loaded onto other computers or processors to implement different aspects. Computer readable media can be non-transitory in some instances.

“The terms ‘program? “Program?” or’software? are generic terms that can be used to refer to any type of computer code or set of instructions. “Software” is used in this context to mean any type of computer code, or set of instructions that can be used to program a computer to execute various aspects of the above. It should also be noted that one or more computer program that executes methods of this disclosure does not have to reside on one computer or processor. They may be distributed modularly among multiple computers or processors in order to implement different aspects of the disclosure.

Computer-executable instructions can be in many forms such as program modules. They are executed by one or more computers, or any other device. Program modules can include routines, programs and objects as well as data structures. Modules that are used to perform specific tasks or implement certain abstract data types. The functionality of program modules can be combined or distributed in different ways depending on the need.

“Data structures can also be stored on any computer-readable media. Data structures can be simplified by showing fields that can be related to their location within the data structure. These relationships can also be established by assigning storage to the fields that correspond with their locations on a computer-readable medium. Any mechanism that can establish a relationship between elements of a data structure may be used, such as pointers, tags, or any other mechanism that establishes relationship between them, could be used.

“The embodiments described above can be implemented in many ways. The embodiments can be implemented in hardware, software, or a combination of both. The software code can be executed on any processor or group of processors. This is true regardless of whether the processors are located in one computer or distributed across multiple computers. Any component or collection that performs the functions mentioned above can be considered a controller. You can implement a controller in many ways. For example, you could use dedicated hardware or general purpose hardware (e.g. one or more processors) to execute the functions. The controller may also be programmed with microcode or software.

“Moreover, it is important to understand that a computer can be embedded in any number of forms such as a rack-mounted, desktop, laptop, or tablet computer. These are just a few examples. A computer can also be embedded in devices that are not considered computers but have suitable processing capabilities. This includes a Personal Digital Assistant (PDA), smartphone, or other portable or fixed electronic device.

A computer can also have input and output devices. These devices can be used to provide a user interface, among other purposes. Printing machines or display screens can be used for visual output, as well as speakers or other sound-generating devices to present audio output. Keyboards and other input devices such as touchpads, digitizing tablets, and mice are all examples of devices that could be used to provide a user interface. Another example is speech recognition, which can be used to receive input information from a computer.

These computers can be connected to one or more networks, in any form, including a local network or a large area network such as an enterprise network and intelligent network (IN), or the Internet. These networks can be built on any technology and can operate according to any protocol. They may include wired networks, wireless networks, or fiber optic networks.

“Also as mentioned, certain aspects can be embodied in one or more methods. You can arrange the acts that make up the method in any way you like. It is possible to create embodiments in which the acts are performed in a different order than what is illustrated. This may allow for simultaneous performance of some acts, even though they are shown in sequential sequence in the illustrative embodiments.

“All definitions as defined and used herein should be understood to control dictionary definitions, definitions within documents incorporated by reference, and/or the ordinary meanings of defined terms.”

“The indefinite articles?a? “The indefinite articles?a?? If not clearly stated, the terms?an,’ and?an,? should be taken to refer to?at most one.

“The expression?and/or? “The phrase?and/or, as it is used in the specification and the claims, should mean?either/both? The elements so conjoined are elements that are conjunctively and disjunctively found in different cases. Multiple elements are listed with?and/or. Multiple elements listed with?and/or should be understood in the same way, i.e.?one? or?more? All elements are so connected. Optionally, other elements can be present than those specifically identified by the “and/or” clause. Regardless of whether they are related or not to the elements identified, other elements may optionally be present. As an example, the reference to “A” and/or “B” can be used with open-ended language like “comprising”. In one embodiment, A can be used to refer to only A (optionally including additional elements than B); in another embodiment to B only (optionally incorporating elements other than A); and in yet another embodiment to both A (optionally containing other elements); and so on.

“As used in the specification and the claims, the phrase “at least one” means at least one element. When referring to a list containing one or more elements, it should be understood that at least one element is selected from any of the elements in this list. However, not all elements are included in the list. This definition allows elements to optionally exist other than those elements that are specifically identified in the list of elements. Refers to elements that are related or not related to the element specifically identified. As an example, let’s say that?at least one? of the elements A or B is missing. Or, alternatively,?at minimum one of A and/or B? Or, equivalently,?at minimum one of A or B? can be used to refer to, in one embodiment to at minimum one, optionally containing more than one A (and optionally incorporating elements other than B); in another embodiment to at most one, optionally involving more than one B (and optionally incorporating elements other than A); in yet another embodiment to at least one (optionally including more that one A) and at least one (optionally including more B); etc.

“Also the terminology and phraseology used in this document are for description purposes only and should not be considered limiting. The use of the word?including? ?comprising,? Or?having? ?containing,? ?involving,? “Integrating” (and variations thereof) is intended to include the items listed below and their equivalents as well as any additional items.

“All transitional phrases, such as?comprising?, are included in the claims and the specification. ?including,? ?carrying,? ?having,? ?containing,? ?involving,? ?holding,? ?composed of,? The like and similar are to be understood as open-ended. This means that they can include but not limit. The transitional phrases ‘consisting of’ and?consisting essentially of are not allowed. Only the transitional phrases?consisting of? shall be either semi-closed or closed transitional phrases, depending on the context.

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