Patent Application
20230249006
The present invention relates generally to methods and devices for transcranial focused ultrasound neuromodulation.
There are a large number of neurological diseases and injuries (too numerous to list here) where neurostimulation is potentially beneficial either to cure the patient or resulting in a significant improvement in the quality of their lives. The essential problem with implementing neurostimulation on a large scale is that the conventional method of achieving neurostimulation requires surgical implantation of electrodes in close proximity to the target volume. There are many possible undesirable complications with surgical implantation of electrodes next to neural tissue, most common is the risk of infection, which could be very serious. Therefore, it is desirable to have a method of neurostimulation that is non-invasive.
Currently, the leading candidate technologies are TMS (Transcranial Magnetic Stimulation) and tDCS (transcranial Direct Current Stimulation). Both methods achieve electrical stimulation non-invasively instead of surgically implanted electrodes. However, the problem with these technologies is that the spatial resolution is very poor (i.e., the target tissue is necessarily quite large) and these methods cannot target tissues deep in the human brain.
An alternative technology is to use low-power focused ultrasound, which can reach any arbitrary target tissue in the brain, regardless of depth. It also has high spatial-temporal resolution compared to the other non-invasive methods. Ultrasonic neurostimulation is a relatively new technology and the mechanisms and optimal stimulus parameters are still unknown. In the published literature scientists typically use a single transducer. Compared to other non-invasive methods, such as transcranial magnetic stimulation (TMS), ultrasonic neuro-stimulation has the advantage that it can reach deep into the brain or other tissue with higher spatial resolution and still maintain high temporal resolution. The conventional approach to achieving ultrasonic neuro-stimulation is to use a single transducer or a transducer array that is focused on the target volume.
An aspect of this work is to optimize the effectiveness (i.e., achieve therapeutic levels of stimulation with minimum power) and spatial specificity (i.e., limit stimulation to small targeted volumes) by creating an interference pattern between at least two transducers. It is a common practice to use multiple transducers in an array to decrease the focal volume, but that is not what we are describing in this invention. Rather, it is the interference pattern between two or more ultrasonic waves that occurs under very specific conditions that achieves more effective stimulation in a smaller volume.
The approach of this work is born from ex vivo research and a radiation force theory of ultrasonic neurostimulation. It strongly suggests using two transducers opposed to one another, such that the interference pattern generated will either produce standing waves (the frequencies are the same) or a beat pattern (the frequencies are slightly different) that will result in much more effective stimulation (less power to achieve a given level of stimulation, or more stimulation for a given power level). It also has the potential to stimulate smaller volumes. These different interference patterns will produce different space-time patterns of mechanical strain arising from radiation force that will alter neurostimulation. For any given medical condition, it is currently unknown exactly what type of neurostimulation is most effective. This approach is intended to optimize ultrasonic neurostimulation by giving clinicians and researchers the flexibility to try different methods of stimulation on the patient to discover what works best for that patient and disease. This will enable translation of this technology from a research novelty to a versatile clinical tool. The device is noninvasive or minimally invasive and can be used for many applications to change the activity of biological tissues including the nervous system.
This approach could be used on any part of the nervous system, central or peripheral, including all sensory systems. It might be used to stimulate, inhibit, or modulate neural activity for the purpose of treating various nervous disorders, too numerous to explicitly list here. There are various ultrasound stimulation parameters such a carrier frequency, pulse repetition frequency, duty cycle, intensity, pulse duration and time in between pulses that can be varied to obtain optimal values for a given clinical condition. In the case of two transducers there are additional parameters that specify the relationship between the two transducers, such as phase. Furthermore, the second transducer may operate at a different carrier frequency from the first transducer and other parameters may be different.
Particularly preferred embodiments provide the capability to switch at will between at least three operating modes: 1) standing wave mode with the two transducers operating at the same frequency, 2) oscillating mode with the two transducers operating at different frequencies, and 3) impulse mode from one of the two transducers only.
There are a multitude of possible medical applications, such as depression, Parkinson's, tremor, dystonia, Epilepsy, pain management, and prosthetic devices for sensory systems. For any condition in which electrical neural stimulation is therapeutic, ultrasonic neurostimulation is a possible non-invasive substitute. This could be in the central or peripheral nervous systems, including sensory systems. Significant advantages are provided. Compared to conventional electrical stimulation, ultrasonic neurostimulation is non-invasive. Non-invasive methods have less risk and cost because no surgery is involved. This makes it much more attractive to patients. The other non-invasive methods of neurostimulation, TMS (transcranial magnetic stimulation) and tDCS (transcranial direct current stimulation) have poor spatial resolution and cannot reach deep into the brain. The conventional method of ultrasonic stimulation, as shown in the literature, uses a single transducer. We use two transducers opposed to one another such that the resulting interference pattern is optimal for stimulation. Here ultrasonic neural stimulation is defined as any combination of stimulation, inhibition and/or modulation of neural activity.
In another aspect, the invention provides an improvement of the above-described work. Two facing focused transducers are used to set up standing waves, then chirp excitation, or other frequency-varying signal, is used to generate radiation force in a smaller spot. Significant advantages are provided. In one particular implementation, we use chirp signals to excite the transducers and gain a reduction in the spot size where the radiation force is applied. Another advantage of using the chirp is the ability to delay the application of the chirp on one transducer with respect to the other and thus scan the location of the stimulation inside the body electronically. Commercial applications include all neuromodulation applications in the brain or other parts of the body where spatial resolution is important and access is available to provide two counter-propagating waves.
In one aspect, the invention provides a method for transcranial focused ultrasound neuromodulation comprising: generating an excitation waveform signal with a controller configured to control operation of the acoustic transducers; exciting two ultrasound transducers using the excitation waveform signal to produce acoustic outputs comprising substantially coincident focus regions resulting in a pattern interference radiation force between the two ultrasound transducers; wherein the excitation waveform signal has a time-varying frequency within range containing a resonance frequency of the two transducers. Preferably, the excitation waveform comprises a chirp waveform. The frequency modulated waveform is not necessarily a chirp waveform, but can more generally be various other kinds of waveforms with variable frequency. It is preferred to have broadband frequency range that varies uniformly. In a preferred embodiment, the chirp waveform comprises a linear chirp waveform. Preferably, the excitation waveform comprises a chirp-up waveform and a chirp-down waveform. The up-chirp followed by a down-chirp, is preferred in order to make the continuous input waveform smoother. For example, if using only up-chirp or only down-chirp, the output frequency would have to change instantaneously when the first chirp waveform ends and the next chirp waveform starts. This could lead to undesirable forms of output pressure, which can also lead to unwanted radiation force field. However, if the ultrasound transducer has a broadband transmit sensitivity and can make the output pressure exactly equal to the input voltage, it will have no problem using only the up-chirp or the down-chirp waveform. In some implementations, the chirp-up and chirp-down waveforms each have 10 cycles. Generally, the number of cycles in the waveform is selected depending on the specific application. The pressure and force field can change based on the number of cycles. If the number of cycles are too low, it can show repeating pattern. However, if the number of cycles are too high, it can be hard to say that the radiation force is an averaged value. Preferably, the number of cycles should be minimal unless you see a repeating pattern.
In one implementation, the time-varying frequency varies within the range 300-700 kHz. The frequency range is not limited to these values but is selected based on various considerations. In general, the wider the interval between the lowest and highest frequencies, the better the results. However, the output pressure is limited due to the transmit sensitivity of the ultrasound transducer. The frequency was determined from the transmit sensitivity of the ultrasonic transducer. Ultrasonic transducers with as wide bandwidth as possible are preferred, as lower frequencies may have lower attenuation in the skull and higher frequencies may result in smaller target sizes. The preferred minimum frequency is 100 kHz, but the wide bandwidth is more important.
In another aspect, the two ultrasound transducers may be excited simultaneously with the excitation waveform, to center the focused radiation force between the transducers, or excited at different times, separated by a controlled delay, to scan the position of the focused radiation force between the transducers. The delay may be implemented using any of various types of time delay logic circuits.
The controller is configured to allow a user of the apparatus to select from at least a first operating mode and a second operating mode. Here the first operating mode (
The second operating mode (
The controller is configured to drive the two or more acoustic transducers independently of each other. This provides electronic control of an acoustic interference pattern generated by the two or more acoustic sources that is provided in the region of the patient.
Preferably some or all of transducers 102, 104, 106 (optional) are acoustic transducer arrays, as schematically shown. Such transducer arrays can be configured to alter a focus position of the acoustic transducer array using phase control of elements of the acoustic transducer array. Such transducer arrays are well known in the art, and so are not further described here.
The apparatus can be configured to modify neural activity in the central nervous system and/or in the peripheral nervous system. Modifying neural activity can include stimulating neural activity, inhibiting neural activity and modulating neural activity in any combination.
Optionally the apparatus can provide a third operating mode that is an impulse mode with only one of the two or more acoustic transducers activated.
Preferably, the two or more acoustic transducers include two diametrically opposed acoustic transducers, as in the examples of
Previously, ultrasonic neurostimulation has been performed without a detailed understanding of the responsible mechanisms. Although stimulation methods have been improved by systematically varying stimulus parameters, in practice, the number of parameters that can be changed is very large. It is not practical to try all possibilities, in all animal models, and in all possible applications, to find the optimal stimulus for any given situation. Typically, ultrasonic neurostimulation stimuli either replicate parameters that have been published, or a few parameters are varied until an improvement is observed. Achieving an optimal stimulus (obtaining the greatest therapeutic effect with minimal delivered energy) requires an understanding of the mechanism by which ultrasonic neuro-stimulation functions. We provide a theory of ultrasonic neuro-stimulation from which we derive an optimal stimulation method.
Ultrasonic acoustic pressure waves oscillate at frequencies too high to have a direct effect on biological tissue. When acoustic wave energy is absorbed or reflected, there is a transfer of momentum that results in a static force known as radiation force, a phenomenon common to all kinds of waves, including light. This radiation force does not oscillate at the acoustic frequency, rather it remains constant so long as the ultrasonic stimulus is on. The strength of this radiation force is determined by the intensity of the acoustic pressure wave and the amount of energy that is absorbed or reflected. It is well known that this force can physically move tissue on the order of microns without harming patients. The physical displacement from radiation force puts stress on tissue and this stress is converted into electrical activity in neurons most likely by mechano-sensitive ion channels that are present in the nervous system. To optimize the efficiency and spatial specificity of neurostimulation the objective is to maximize stress caused by radiation force in the smallest volume. Experimental evidence in support of this theory is described in detail in section B below.
It is known that acoustic resonators can be built that will exert mechanical forces on small objects such as single cells to manipulate them without harming them. This field is known as acoustofluidics. It is a useful technology to manipulate single cells into a particular location for some kind of measurement. These resonators are built with specific dimensions matched to the acoustic frequency being used to establish standing waves. These standing waves are created either with single transducer and a reflector or by having two opposing transducers (
The in vitro experimental condition does not mimic in vivo clinical conditions. The presence of the rigid MEA has a profound effect on our results and would not be present in vivo. However, we can duplicate the effect of the MEA in vivo by either using either a reflective element or a second transducer on the opposite side of the target tissue to make a resonant cavity and generate standing waves as shown conceptually in
This approach can be used on any part of the nervous system, central or peripheral, including all sensory systems. It might be used to stimulate, inhibit or modulate neural activity for the purpose of treating various nervous disorders, too numerous to explicitly list here. There are various ultrasound stimulation parameters such as carrier frequency, pulse repetition frequency, intensity, and pulse duration that might be varied to obtain optimal values for a given clinical condition. In the case of two transducers there are additional parameters that specify the relationship between the two transducers, such as phase. Furthermore, the second transducer may operate at a different carrier frequency from the first transducer and other parameters may be different. Although it is generally the case that the focal volume corresponds to the targeted neural tissue, that is not the case in version #3 described below.
Exemplary embodiments include:
1) Standing wave version (
2) Dynamic radiation force version (
3) Offset focal volumes version (
4) Space-time dynamic radiation force versions (applicable to any of the above versions). Instead of having two temporally static carrier frequencies, we can vary both frequencies (substitute f1(t) for f1 and f2(t) for f2 in the above versions) in time to alter the radiation force in both space and time electronically. The temporal variation of frequency for the two transducers might be the same such that the difference frequency remains the same in time, or it might be different, so that the difference frequency itself varies with time. Likewise, the phases of the two carrier frequencies can be dynamically altered either to maintain a similar phase difference or to create a dynamically changing phase difference. This serves a similar function as changing the frequencies by permitting electronically controlled changes to radiation force in space and time.
For the results of
Ultrasonic neuromodulation has been demonstrated in the brains of human, monkey, sheep, rat, mouse, and retina of salamander and rat. The capability of ultrasound to reach any brain structure noninvasively through the skull, and the highly developed technology to deliver ultrasound make this approach promising both for basic studies of neural function and clinical applications. Yet results in different preparations have varied, including both excitatory and inhibitory effects. The development of this approach would benefit greatly from a quantitative understanding of the mechanisms of ultrasonic neuromodulation, allowing the process to be optimized in terms of efficacy of stimuli, efficiency and spatiotemporal distribution of effects.
In the process of transduction of a stimulus into a biological response, one can distinguish the physical mechanism such as acoustic pressure or thermal energy from the biophysical mechanism that senses that energy, including changes in membrane capacitance or particular ionic channels. Here we focus on the physical mechanism by which an acoustic wave is converted into an effective stimulus for a neuron, a process that is currently not understood. The leading candidates for physical mechanism are radiation pressure, the process by which an absorbed or reflected wave delivers momentum, and cavitation, which includes the stable or unstable formation of bubbles, creating a mechanical disturbance, and thermal energy.
Radiation force is a nonlinear effect proportional to the acoustic wave amplitude, thus creating a continuous, non-oscillating force for a stimulus of constant amplitude. By this mechanism a carrier wave with a frequency too high to have a direct biological effect can be converted into a low frequency mechanical force with dynamics of the envelope of the wave. When radiation force is exerted on a liquid, this results in bulk flow of fluid known as acoustic streaming. Tissue attenuation increases with carrier frequency, therefore radiation force and also heating will increase with frequency.
Cavitation can occur if the acoustic pressure wave becomes sufficiently negative, causing gas bubbles to form that oscillate at the carrier frequency. Inertial cavitation occurs when those oscillations change in size and eventually burst the bubble, creating a destructive violent event. In stable cavitation the bubble does not burst and is hypothesized to produce safe neuromodulation. Cavitation is less likely at higher carrier frequencies because it becomes more difficult to sustain oscillations in the bubble.
In this study we use optical imaging to measure displacements in the retina and vary the acoustic frequency to test which of these mechanisms is most likely. We find that ultrasonic stimulation in the retina is consistent with a model whereby radiation force produced micron-scale mechanical displacements. The acoustic frequency dependence is consistent with radiation force but inconsistent with cavitation. In addition, we see that standing waves influence the effects of ultrasound. We conclude that radiation force is the primary physical mechanism for ultrasound to stimulate the retina.
We conducted several different types of experiments for the purpose of establishing whether radiation force is the dominant physical mechanism responsible for ultrasonic neurostimulation, the details of these experiments are described below in the appropriate sub-sections. In terms of experimental design, our goal is to use at least two retinas from two different salamanders and record from at least 30 ganglion cells for electrophysiology experiments. For the imaging experiments we used three retinas from three salamanders, for the 43 MHz electrophysiology we used three retinas from three salamanders, for 15 MHz recordings we used two retinas from two salamanders, and for lower frequency recordings (2.9 MHz, 1.9 MHz and 500 kHz) we used two retinas from two salamanders. For electrophysiological recordings the number of cells for each experiment is indicated either in the figures or in the figure legends. Data is available upon request.
All statistics used have well known definitions and require no additional information (e.g., mean, standard error of the mean, squared Pearson correlation coefficient r2).
Multielectrode array (MEA) recordings were performed as described in the literature. The isolated retina of the tiger salamander of either sex was adhered by surface tension to a dialysis membrane (Spectrapor 7 50000, Fisher Scientific) attached to a custom Delrin holder. The holder was placed on a motorized micromanipulator (MP-385-2, Sutter) and lowered onto a multielectrode electrode array (ThinMEA, Multichannel Systems) ganglion cell side down. For 43 MHz experiments where the focal spot<100 μm, a high-density array was used (5×6, 10 μm diameter electrode, 30 μm spacing). For all other lower frequency experiments, a lower density array was used (8×8, 10 μm dia., 100 μm spacing), which better matches the focal spot size. Full field flashes from a red LED were sometimes used to verify that ganglion cells were responding normally to visual stimuli, especially if conditions of ultrasound stimulation did not show a response. Error bars are SEM unless otherwise noted.
We used four different transducers, 43 MHz (custom), 15 MHz (Panametrics, A3195, 0.5″ dia., 2″ FL), 2.25 MHz (Olympus, V305) and 0.5 MHz (Olympus, V301) in order to span a large frequency range. The 2.25 MHz transducer had a relatively wide bandwidth and was operated at multiple frequencies (1.9 and 2.9 MHz). Transducers (15, 2.25 and 0.5 MHz) were fitted with a water-filled cone that was sealed with either parafilm (2.25 and 0.5 MHz) or plastic wrap (15 MHz) and mounted on a motorized micromanipulator (MP-385-2, Sutter). A camera from below was used to position the transducers so that the center of the focal spot was in the center of the array. Transducers were lowered into the bath above the retina, and height was adjusted so that the focal point was on the retina. Ultrasound propagated from the transducer, through the water-filled cone, perfusion fluid, dialysis membrane, retina, and then reflected off the glass/metal surface of the MEA (3A on
The styryl dye FM4-64 was bath applied by immersing the isolated retina in a concentration of 82 μM (100 μg in 2 ml) FM4-64 in oxygenated Ringer's for one hour prior to placement on the MEA. This dye inserts itself in the outer leaflet of the cell membrane where it becomes fluorescent, allowing us to image changes in position and shape of the cell membrane with ultrasonic stimulation.
A custom two-photon laser scanning microscope in the inverted configuration was used to image the retina during ultrasonic stimulation. A simplified diagram is shown in 3A of
The mirror position of the scanning galvanometers was recorded on the same computer that generates and records the ultrasound stimuli (one second on, one second off, 40 W/cm2), allowing us to compute the timing of the image at any pixel relative to the ultrasound stimulus. The laser scanning and ultrasound stimulus were desynchronized, such that any given pixel will be recorded at random times during the two second period of ultrasound stimulation. In theory we can get temporal resolution approaching the dwell time of a single pixel with a sufficiently large data set; in practice we binned the data in 10 ms bins for a sufficient signal-to-noise ratio. To compute vector fields reflecting the effect of ultrasound, we used unwarpJ, which is an imageJ plug-in that performs spline-based elastic registration of two images. We compared steady state images in the ultrasound on and off conditions.
Because of the large area of scanning at a high frame rate (20 Hz), we corrected for most distortion at the edge of the frame by computing the average actual mirror position based on control experiments recording mirror positions with slow mirror velocities and no distortion and then recording actual mirror positions at high velocities. This distortion does not affect our analysis, which is based on changes in the images as a function of time in the cycle of ultrasonic stimulation.
To model the observed displacement as an effect of radiation force, we assumed the ultrasound field at 43 MHz was transmitted through a multilayered medium composed of water, retina, glass, and air, and calculated assuming 40 W/cm2 incident power. The retina layer was 150 μm thick and the glass layer was 180 μm thick. Water and air media were assumed to be half-spaces. The model includes the following: acoustic streaming (the bulk flow of fluid from radiation force acting on the liquid medium between the transducer and the retina), radiation pressure on the retina-water and retina-glass interfaces, radiation pressure from absorption in the retina, and the interference pattern that is a consequence of wave reflection off the MEA. The speed of sound and density parameters are standard values from the literature, the attenuation at 43 MHz was set at 4.01 dB/cm for water and 22.02 dB/cm for retina. Estimated radiation pressures were then used in COMSOL Multiphysics finite element software to calculate the deformation of the retina in response to ultrasound. The retina was considered to be an incompressible material (i.e., Poisson ratio of 0.5). We determined the value of the elastic modulus (0.5 kPa) that gives 4 μm of maximum displacement in the direction of the propagating wave as seen in the data. Since soft tissues such as retina exhibit large deformation with nonlinear strains, a large deformation model was used to estimate the displacement field in the retina.
A COMSOL model to simulate standing waves was similar to the model described above except that after the incident wave reflected off the MEA, we allowed for that wave to perfectly reflect off the face of the transducer. In this case, retinal deformation was not calculated. The frequency is set to 1.9 MHz and the nominal intensity is 56 W/cm2 (intensity at the focus in free space). Standing wave amplitudes were calculated by averaging intensity values on-axis in the retina. Zero distance refers to the reference position where the focal point is coincident with the retina-MEA interface.
A quantitative radiation force model was used to fit neural population activity in the retina. The model is based on analytic equations valid for linear low-amplitude ultrasound in free space from the literature. This model does not account for the reflection off of the MEA and the resulting standing waves, the presence of coupling cones, and the dialysis membrane. For 1.9 MHz, where standing wave effects are large, we use the data from the transducer position with the lowest threshold. The analytic expression takes as input the absorption coefficient (retina is similar to brain, so we used a brain absorption coefficient value from the literature), the carrier frequency (f), intensity (I), radius of the transducer (a), and focal length (d), to estimate radiation force in three-dimensional space expressed in a cylindrical coordinate system where x is axial distance from the transducer and r is the radial distance away from the central axis. Radiation force weighted unit volume (i.e., Σw(x,r)RF(x,r), 10 μm grid in cylindrical coordinates, approximately the size of cell somas) is summed over a volume defined by three parameters: between two radii r1 and r2 and depth, and then passed through a sigmoid. The optimal values for salamander retina were determined to be the smallest possible volume around the focus r1=0, r2=5 μm, depth=10 μm, equivalent to the maximum radiation force. (
Ultrasonic stimulation was applied to a preparation used for measuring neural activity consisting of the retina placed on a planar multielectrode array (MEA) patterned on a glass slide (180 μm thickness). We imaged the retina through the glass MEA with a two-photon laser-scanning microscope after applying the membrane dye FM 4-64 to the bathing medium (3A on
We converted these displacements between steady-state On and Off ultrasound into a vector field using image processing software (see Methods) (
To interpret the potential mechanism of this displacement, we modeled the expected mechanical response of the retina from radiation force using finite element analysis (COMSOL). In the simulation we considered the following factors: acoustic streaming (the bulk flow of fluid due to radiation force acting on a fluid), reflection from the water-retina and retina-MEA interfaces, absorption in the retina, and the interference pattern that results from the wave reflecting off the MEA. It was determined that 88% of displacement can be accounted for by the combination of reflection from the water-retina interface combined with the interference pattern (i.e., standing wave) that comes from wave reflection off the MEA, 9% comes from absorption, and 3% from acoustic streaming. A key parameter for this calculation is the Young's modulus of elasticity for the retina. However, the literature has values that vary by three orders of magnitude, depending on the method of measurement. We thus allowed the Young's modulus to be a free parameter and fit the model to account for the maximum observed displacement, which was 4 microns. The resulting value of Young's modulus was 0.5 kPa, which is close to the range found in the retina (0.1-2.0 kPa) with the scanning force microscopy method. The general features of the simulated vector field qualitatively matched the experimental vector field of displacement: large downward motion in the outer retina right under the focus which decreases to zero at the level of the MEA (
To quantify the displacement, we found a region with high image contrast with the largest displacement and examined the change of displacement in 10 ms time bins (
The vertical displacement occurred very rapidly (<10 ms) consistent with the expected temporal dynamics of radiation force. The fast onset of displacement is consistent with the fast response of neurons to ultrasonic stimulation. The recovery to baseline, which reflected the elastic properties of the retina was slower and was fit by double exponential with time constants of 33 ms and 530 ms (
To examine the relationship between displacement and neural activity, we then compared measurements of these two quantities as a function of stimulus intensity. We imaged a level in the retina above the IPL midway through the retina that showed considerable lateral displacement. We varied the ultrasound intensity from below the threshold of neural activation to above the level of a saturating response (
Absorption increases with higher acoustic frequency, and thus both radiation force and heating are expected to increase with higher carrier frequency. In contrast, the probability of cavitation decreases with higher carrier frequency because of the shorter time interval available to cause a bubble to form out of solution and to keep it oscillating. Many protocols of ultrasonic neurostimulation use lower frequencies (<1 MHz) to allow sufficient energy to penetrate the skull, which is known as transcranial neurostimulation. It is conceivable that at lower frequencies a different mechanism such as cavitation is involved. We therefore changed carrier frequency in several steps between 43 MHz and 0.5 MHz to measure activation of retinal ganglion cells by ultrasound at different frequencies on the retina.
To more completely characterize the response at a given frequency, we varied both pulse intensity and duration across a wide range for the 43 MHz transducer (
Using the same stimuli, at 15 MHz a greater intensity was required to stimulate neurons compared to 43 MHz (
We found that at 43, 15 and 1.9 MHz, the intensity at half maximal varied as a function of frequency raised to a power of 1.27 (
We then tested whether retinal neural activity could be fit with a single quantitative model of radiation force across the range of intensities and frequencies tested. The model is structured to be the simplest possible that minimizes RMS error between the model and data. The neural response was assumed to be proportional to the sum of radiation force over some unknown volume followed by a sigmoidal non-linearity. We used an analytical model of radiation force valid for linear low-amplitude ultrasound in free space, which has absorption coefficient, the carrier frequency, intensity, radius of the transducer, and focal length as parameters, to estimate radiation force in three-dimensional space expressed in a cylindrical coordinate system. From this model, we computed the radiation force for each intensity, transducer and spatial location, and then passed this value through a stage of spatial integration representing the neural properties that sense ultrasound and then a sigmoidal function to predict neural activity. The free parameters of the model defined the volume of spatial integration and shape of the sigmoid, which was fixed across all intensities and acoustic frequencies. In the retina, the spatial integration was centered on the transducer focus, and the optimized scale of integration was small (10 μm diameter, 10 μm depth), which was equivalent to computing the maximum radiation force. This model showed that the analytically computed maximum radiation force could be used to predict the neural response from 1.9 to 43 MHz with a single sigmoidal neural activation function in the retina (
We measured temperature rise under our experimental conditions using small (76 μm) thermocouples (J and K type, OMEGA) placed on the array with a retina held in place on top of the thermocouple. With the perfusion running as during ultrasonic stimulation, the temperature change is not measurable at 60 W/cm2 and 15 MHz, and without perfusion, we measure only 0.1-0.2° C. increase. Small thermocouples suffer from sources of artifact such as ultrasound reflection off the thermocouple and conduction of heat away from the source and heating due to friction from radiation force moving the transducer relative to tissue. This latter artifact is much greater with pulsed ultrasound, where the thermocouple will oscillate at the pulse repetition frequency, whereas we are using continuous wave. We expect that these artifacts are minor as the thermocouple was attached to the bottom of the dish thereby reducing friction since the thermocouple cannot move. The ultrasound will reflect off the glass surface of the MEA in any event, so reflection off the thermocouple is not significantly different from the normal experimental condition. We found that perfusion removed heat much more effectively than the thermocouple wire, such that under normal conditions of ultrasound stimulation with perfusion running, we cannot measure any temperature rise from ultrasound. Although these studies do not categorically rule out thermal effects at a fine spatial scale, we find no evidence of significant thermal effects. Higher resolution spatial-temporal measurements of temperature changes in the future will be useful to examine if any thermal effects do exist.
In the retinal preparation, below the tissue is a glass MEA of thickness 180 μm, followed by an air space. The top and bottom surfaces of the MEA create a large mismatch in acoustic impedance, which is expected to reflect ultrasound. Thus, the space between the transducer and MEA may form a cavity that could generate a standing wave, where locations spaced at one-half the acoustic wavelength (λ) would experience destructive interference (nodes), and intervening locations experiencing constructive interference (anti-nodes). The acoustic pressure in the standing wave is converted into radiation pressure through absorption generating alternating high and low pressure volumes that do not temporally modulate at the carrier frequency. The relationship between acoustic pressure standing waves and radiation pressure is well known in micro-fluidics, where it is used to physically move small particles, including individual biological cells, to a desired location. Radiation pressure is greatest at anti-nodes and smallest at nodes, causing tissue at nodes to be compressed by adjacent high pressure anti-nodes and tissue at anti-nodes to be stretched by adjacent low pressure nodes. Such mechanical pressure on tissue could have an additional influence on neural activity. We tested the neural effects of standing waves by simply changing the distance between the transducer and the MEA. This will not change the locations of the nodes and anti-nodes as they are fixed by the carrier frequency, but the change in cavity length will affect the amplitude of standing waves, with a maximal standing wave amplitude when the cavity length is a multiple of λ/2. To illustrate this effect, we computed a COMSOL simulation of the transducer-electrode array cavity with a perfect reflection off the transducer face, which produced a large modulation in acoustic intensity in the retina with a characteristic period of λ/2 in transducer distance (
We tested the effects of standing waves at relatively low frequencies, 2.9 MHz (λ=517 μm, λ/4=129 μm) and 1.9 MHz (λ=789 μm, λ/4=197 μm), close to where most ultrasonic neurostimulation studies are conducted, yet high enough that we can still get robust responses, and where the λ/4 distance is large and comparable to the thickness of the retina (˜120-150 μm). The ultrasound stimulus was a continuous wave 100 ms pulse, which we had previously found to be optimal at higher frequencies (
We found that the firing rate of some cells was very strongly modulated by the distance between the transducer and the MEA with a period of λ/2, consistent with standing waves (
We then tested whether standing waves were necessary for neurostimulation by tilting the transducer at an angle of 27° to vertical. Although a spatial interference pattern still occurs between the incident and reflected waves, the depth of modulation will not be as great as when the transducer is positioned vertically, and such a pattern would move with distance between the transducer and glass. We found that the tilted transducer condition still generated a response (
Our results show that ultrasonic neurostimulation in the retina produces radiation force and micron-scale displacement. A quantitative model of radiation force across multiple acoustic frequencies and power levels indicates that radiation force is the likely physical mechanism of action. We further show that standing waves can modulate neural activity, suggesting a potential new method to further control activity.
Estimates of temperature rise based on ISPTA, pulse duration, density and specific heat capacity and absorption coefficient are very small (0.007° C.-0.04° C. for in vivo studies in humans and sheep. Furthermore, these methods assume all energy goes to an increase in temperature, and do not account for heat loss by conduction or convection, so the actual temperature rise should be lower. Previously, at 43 MHz and 30 W/cm2, well above stimulation threshold, we could not measure a temperature rise with the perfusion running, although we could measure a 0.5 deg. C. increase from prolonged stimulation without perfusion. In a literature study using C. elegans, mutants lacking thermosensitive receptors behaved like wild type animals, while mutants that lack touch sensory neurons have an impaired response to ultrasound. Together, there is no evidence for heating as a physical mechanism for brief ultrasonic neurostimulation.
We found using the same transducers, amplifier, frequencies and power settings that successfully stimulated in vivo mouse that in the retina, higher acoustic frequencies were more effective than lower frequencies, thus ruling out cavitation as a possible physical mechanism (
A hypothesis of ultrasonic neurostimulation is neuronal intramembrane cavitation excitation (NICE), which is a theoretical model that has been fit to empirical results. The intramembrane cavitation hypothesis asserts that stable cavitation exists inside the cell membrane causing a change in cell capacitance that ultimately leads to action potential firing. Although this model has been fit to various in vivo experimental data, it does not describe our data because of the strong correlation between greater neural activity and higher ultrasonic frequency.
It has been known since ancient Greece that mechanical deformation of the eyeball generates pressure phosphenes (the appearance of light when there is none). Although it is still not known which retinal cells are responsible, it is clear that mechanical force can result in ganglion cell activity. Studies with deformation of the cat eyeball showed that different ganglion cells respond through network stimulation, likely in the outer retina. Most importantly they conclude that mechanical strain is the cause, not retinal ischemia from high intra-ocular pressure. The authors speculated that inhibitory horizontal cells or amacrine cells, might be sensitive to strain because of their lateral connections. A phenomenon that has been known for thousands of years supports the concept of mechanical strain on neurons as the cause for this neural stimulation.
Leading candidates for biophysical mechanisms are mechanosensitive ion channels, capacitive effects from mechanical deformation of the cell membrane, and direct effects on endocytosis/exocytosis. A simple biophysical mechanism that could transduce mechanical strain is a change in membrane capacitance, which can result from radiation force. Alternatively, stretching, compressing or bending of the cell membrane may cause the opening or closing of mechanosensitive ion channels, which are found in all parts of the nervous system. These serve different functions such as controlling osmotic pressure to guiding developing neurons. Sensitive channels that are good candidates to convert mechanical stress from ultrasound into neural activity include Piezo, TRAAK, TREK-1, and TREK-2. In a study expressing mechanosensitive ion channels in Xenopus oocyte, ultrasound was found to significantly influence membrane current of the potassium channels and had a small effect on the sodium channel. In C. elegans, ultrasonic neurostimulation requires mechanosensitive channels.
It is known that high static pressure will suppress synaptic activity. This is the physiological basis for High Pressure Neurological Syndrome (HPNS), a danger for deep-sea divers exposed to pressures greater than 1 MPa, but one from which divers fully recover without permanent damage. Potential mechanisms are differential activation of calcium channels, or a direct effect on exocytosis. In general, multiple mechanisms of ultrasonic neurostimulation could operate under different conditions, including stimulus parameters or type of tissue.
A recent approach to modulating neural activity is the genetic targeting of molecules sensitive to ultrasonic stimulation. Termed either ‘sonogenetics’ or ‘acoustic mechanogenetics’, such methods promise to create an alternative to optogenetic approaches that benefit from the depth of penetration possible with ultrasound as compared to light. Key to the design of such approaches is the knowledge of the physical mechanisms by which ultrasound can act. Ultrasound effects acting through radiation force as we have identified here could potentially be used to activate sonogenetic probes. In doing so, in any given tissue, mechanical sensitivity of sonogenetic probes should exceed the endogenous sensitivity to ultrasound.
The large displacements in the outer retina shown in this study are consistent with ultrasonic activation of neurons in the retinal network, which provide inputs to the ganglion cells. Previously we observed that blocking synaptic transmission with CdCl2 abolished ultrasonic neurostimulation, indicating that we were not directly stimulating ganglion cells. One might assume, therefore, that the biophysical mechanisms of transduction are not present in the ganglion cell soma or dendrites. However, our present results show that little displacement was observed in the ganglion cell layer (
Comparison with Another Ex Vivo Retina Study
Another ex vivo retina study at 2.25 MHz in rats showed that neurons frequently exhibit multiple response peaks with a temporal pattern that varies with intensity. We have seen similar effects using 1.9 MHz. Our intensity levels at 1.9 MHz are about one order of magnitude greater those used in this other work which could be due to species differences in either mechanical properties or biophysical mechanisms. The rat retina contains blood vessels making it mechanically stiffer, and salamander retina somas are relatively large (15 μm dia.). Or there may be a different distribution of mechanosensitive ion channels.
Some kind of measurement technology is needed to verify where the ultrasound focus is actually located in the brain and what area of the brain is activated in order to make in vivo behavioral results interpretable. How likely is it that the activated brain region is off-target for in vivo rodent work? When determining the spatial location of the transducer that produces the highest probability of a behavioral response, large variability in this spatial map was found across mice when tested at relatively higher frequencies. This spatial variability most likely arises from the ultrasound focus targeting different areas of the brain in different mice, even when the external location of the transducer is intended to be the same.
In addition, the interpretation of some in vivo results is potentially confounded by the possibility of inadvertent stimulation of the auditory pathway when ultrasound is modulated in the range of audible frequencies. This modulation in the audible frequency range is not present if continuous wave (CW) stimuli are used. The challenge for future studies is to find that region of ultrasound stimulus parameter space that generates direct neurostimulation without the auditory confound. Our studies of the ex vivo retina are of course free from auditory effects, thus making it a useful system to study the parameter space of direct neurostimulation. Our exploration of the stimulus parameter space has shown that a CW pulse of relatively long duration (e.g., 100 ms,
One prior study used CW pulses of 10 ms or shorter, and although another study used 80 ms pulses, they were not CW (a pulse repetition frequency of 1.5 kHz is in the audible range). Neither study used longer duration CW pulses as we have done here. The difficulty in exploring such a large parameter space is a strong motivating factor to develop a theory of ultrasonic neurostimulation that can guide us to the optimal stimulus. One can view stimulation of auditory or vestibular organs as being an off-focus activated brain region where low frequencies will be more effective due to their larger focal volumes. Recent research has demonstrated how to avoid auditory stimulation by smoothing the sharp edges on an 80 ms CW pulse of ultrasound thereby eliminating the auditory brainstem response without affecting motor responses in normal hearing mice.
A number of studies have looked at standing waves in rodent skulls, by both measurement and simulation. The small skull size combined with sub MHz frequencies that are not well focused in the axial direction will generate standing waves as a consequence of reflections off the opposite side of the head. These uncontrolled standing waves have the potential to stimulate parts of the brain that were not targeted.
The conventional approach to ultrasonic neurostimulation is to use a single transducer with a focus at the target (
One ultrasonic neuromodulation study used two confocal transducers operated at 1.75 MHz and 2.25 MHz to generate a beat pattern at 500 kHz. However, a 500 kHz difference frequency is likely too large to be optimally effective. In comparison, vibro-acoustography generally uses difference frequencies in the range of 20 to 100 kHz to mechanically vibrate tissue via dynamic radiation force to measure elasticity.
Many published studies have claimed that “pulsatile” ultrasound (i.e., modulating the carrier with a square wave) is more effective stimulus than continuous wave, often with a PRF in the 1-2 kHz range, but no theoretical reason has been given to explain why. Studies showing that ultrasound can act through an auditory pathway also show that 1 kHz PRF was most effective. There is at least one report of normal human subjects with earplugs inserted hearing high-frequency tones that are correlated with the intensity and pulse repetition frequency of transcranial Doppler ultrasonic imaging. Although it is still unclear how activation of the auditory system might influence the many behavioral or neural recordings, our results suggest that radiation force would translate the envelope of stimulation, i.e., the pulse repetition frequency into a mechanical stimulus. The direct conversion of the envelope of the stimulus to a mechanical stimulus through radiation force could explain why pulse repetition frequencies in the audible range are chosen as being more effective. In vivo researchers should address this potential confound, such as by using deafened animals, using continuous stimuli, or smoothing the edges of ultrasound pulses to reduce audible frequency components. Other potential confounds to direct brain stimulation besides the auditory system are activation of the vestibular and somatosensory systems. Direct measurements of brain activity will be important as opposed to relying exclusively on behavioral outcomes.
There exists a strong theoretical and empirical understanding of using radiation force to exert mechanical strain in the fields of acoustofluidics and elasticity imaging. Here we show a new application of these principles to ultrasonic neurostimulation. Our findings suggest that future approaches to ultrasonic neurostimulation should explore the parameter space defined by these alternative methods of generating radiation force. An understanding of the physical mechanism of action will allow studies in this area to pursue how radiation force might be manipulated to optimize stimulation in different applications and simultaneously provide insights into biophysical mechanisms.
In further developments, the invention provides a technique for transcranial focused ultrasound neuromodulation with improvements in focus, resolution, and scanning capability. As ultrasound can penetrate deep areas, it is suitable for non-invasive neuromodulation. In order to modulate a small area, most current ultrasound neuromodulation uses a single focused transducer. We developed a high spatial neuromodulation method reducing the size of the modulation area by using general focused transducers compared to the case of using single focused transducers. The benefits of our approach are in achieving superior targeting specificity and being able to change the position of the modulation spot electrically rather than with a mechanical method.
In contrast with the present invention, most of the neuromodulation studies using ultrasound use a single transducer. In some cases, two or more transducers are used for higher spatial resolution. However, these cases are not different from making one large size focused transducer with a smaller f-number in the end. For stronger modulation, higher pressure should be applied, and to increase resolution, the size of the total transducer should be increased having a smaller f-number. When high pressure is applied, high temperature rises and tissue damage due to cavitation can occur. Therefore, a method that can perform powerful neuromodulation even with lower pressure is needed.
As described above, a method of maximizing the radiation force by generating a standing wave using two or more focused transducers is provided. In further developments of this method, the invention provides techniques for improve spatial resolution. In addition, the present invention provides techniques for scanning the focus without the need for mechanical movement. In the case of most prior neuromodulation using tFUS, since the size and structure of the transducer are determined, the position of the target or the transducer moves mechanically in order to modulate a different area. To electrically change the target position of the transducer, a transducer array capable of beamforming must be used. However, if the focal point is changed through beamforming, it is difficult to achieve the same performance at all locations as the size and pressure of the focal spot change depending on the location.
ARF used in neuromodulation is a Reynold stress force (RSF), which is widely used in acoustic streaming. It is generated by the second-order term of the wave equation and can be expressed as a first-order pressure field. The RSF can be expressed as the following:
F=−ρ
0
u
1
∇·u
1
+u
1
·∇u
1
where ρ0 is the density and u1 is the velocity. Since ARF is proportional to the gradient of the square of the first-order pressure, when two transducers are placed facing each other to create a standing wave, a stronger radiation force can be obtained compared to a single transducer, which is called PIRF.
More specifically,
We provide a neuromodulation system that improves PIRF by using a chirp waveform instead of a continuous waveform (CW). PIRF using CW generated a stronger radiation force than when using a single transducer, but the focal spot size in the axial direction could not be reduced. However, when a chirp waveform is used instead of CW, the size in the axial direction of the area where the radiation force is applied can be remarkably reduced. In addition, it is possible to move the area where the radiation force is applied in the axial direction in a simple way by giving the one of the transducers a time delay. We developed a system that can dramatically reduce the spot size of neuromodulation using PIRF and a method that can move the spot position quickly and easily.
According to an embodiment of this technique for neuromodulation using chirp waveform pattern interference radiation force, two identical focused transducers are used for PIRF.
According to embodiments of the present invention, if a chirp waveform whose frequency changes with time is generated instead of a single frequency waveform, an ideal standing wave generated over the focal spot is generated only at the focal point. The frequency range of the chirp waveform is selected to give a sufficient change (in our case, +−40%) centered on the resonance frequency (in our case, 500 kHz) of the transducer.
Furthermore, the location where PIRF occurs can be changed by an electrical method rather than a mechanical method. By suitable adjustment of a waveform excitation controller, a time delay is applied to the chirp waveform of one focused transducer relative to the other. As a result, the focus position can be shifted by half the value multiplied by the speed of sound in the medium. For example,
As this approach uses PIRF, neuromodulation is possible with strong radiation force compared to peak pressure. Since it is possible to proceed at a lower pressure than the conventional method, it is possible to reduce tissue damage due to high temperature rise and cavitation. Compared to PIRF using a single frequency continuous wave, when the chirp waveform is used, superior spatial resolution can be secured without changing the characteristics or arrangement of the transducer, and there is no need for frequency tuning to reduce the effects of internal reflection and refraction. Finally, it is possible to quickly move the position of the neuromodulation spot by an electrical method, and precise neuromodulation is possible because the position can be corrected in exact proportion to the time delay.
It is noted that, in general, an ultrasound transducer has a resonance frequency corresponding to it. If it is not driven with a frequency sufficiently close to the resonance frequency, energy efficiency is not maximal. Therefore, if a chirp waveform using multiple frequency ranges is used, energy loss occurs because it does not have the same output at all frequencies. Therefore, it is preferred to avoid using a transducer with a narrow frequency bandwidth, because the output pressure is highly likely not to come out as intended even if a chirp waveform input voltage is applied. Thus, preferably transducers with broad frequency bandwidth are used with this technique.
In one illustrative experiment, we present a method that calculates PIRF with a pressure field and can measure the relative size using a commercial focused ultrasound transducer. The method includes three steps: pressure measurement, polydimethylsiloxane (PDMS) membrane movement measurement and radiation force comparison. The focused transducer used for measurement is a commercial focused transducer with the same structure and characteristics. The transducer has a resonance frequency of 500 kHz, an F-number of 1.25 and a diameter of 25.4 mm. A stick-type hydrophone was used to measure the pressure, and the theoretical radiation force was calculated from the measured pressure.
To measure the radiation force, a very thin PDMS sheet was fabricated and used. The thickness of the PDMS sheet is about 200 μm, and to minimize the effect of force due to ultrasonic reflection and scattering, a material having similar acoustic impedance to water is used. Since the movement of PDMS was smaller than 100 μm, it was recorded using a microscope equipped with a camera. To accurately compare the pressure measurement and the PDMS measurement, the measurement was made at the same location, so the experiment was conducted by making a holder holding the hydrophone and the PDMS sheet at the same time. To accurately grasp the power state of the transducer in the movement of the PDMS recorded with the camera, the measurement was carried out with the laser light on the PDMS. Since the visible area of the microscope was about 3 mm, when the PDMS measurement position was moved, the microscope was also attached to the mechanical stage and recorded as shown in the experimental setup of
Each time the transducer 1512 was turned on and off, sufficient time was given for the PDMS 1514 to stabilize after it started to move. A total of 6 seconds of turning on for 4 seconds and turning off for 2 seconds at each position was recorded for 18 seconds. In the recorded image, it was measured how many pixels the position of the end of the PDMS moved as shown in
Finally, a time delay was applied to the left and right transducers to move the radiation force to the left and right, respectively. Since the speed of sound in water is about 1480 m/s, if a time delay of 4 μs is applied to the left or right transducer, it will move about 1480×4/2=3 mm in that direction, and if a time delay of 8 μs is given, it will move in that direction. It moves twice, about 6 mm. As shown in
There are two major advantages of PIRF, and there are two more advantages when using the chirp waveform. First, compared to the conventional neuromodulation method using a single focused ultrasound transducer, the PIRF neuromodulation method using two transducers generates a greater radiation force at the same acoustic pressure. Because the same performance can be achieved even with a lower pressure, damage caused by heat and cavitation can be reduced. Second, the radiation force generated by a single transducer is a pushing force in one direction, but PIRF is a squeezing force. It is expected that the squeezing force will stimulate more reliably than the radiation force pushing in only one direction. If a force is applied in only one direction, the entire tissue may move, and stimulation may not work well at the desired point. On the other hand, since PIRF squeezes the target tissue from both sides at the same time, there is a high probability that stimulation will occur reliably.
Third, the PIRF with chirp waveform has the advantage of being able to apply a squeezing force to the smallest area compared to all the existing ultrasound neuromodulation methods. All tFUS neuromodulation methods are trying on reducing the size of the stimulating spot. But all those methods could not stimulate the area smaller than the size of the focal spot, which is defined by the wavelength and f-number of the transducer. The PIRF with chirp waveform method overcomes the limitation of all existing tFUS neuromodulation methods, enabling stimulation on an area smaller than the focal spot with a squeezing force.
Finally, it is able to move the modulation spot electrically, which was only possible with a multi-channel array transducer. By controlling the delay time of one of the two transducers, the position of the neuromodulation spot can be moved precisely. The biggest advantage of this technology is that it can be directly applied to the existing ultrasound transducers used for tFUS neuromodulation simply by changing the input signal into a chirp waveform.
Although this description focuses on a technology related to radiation force control for neuromodulation, it can be used in all fields using ARF. Typical applications using ARF include particle manipulation and acoustic tweezers. Both use the ARF generated in the pressure field to move or hold the particle. If the chirp waveform and standing wave are used, the spatial resolution is improved compared to the conventional method, so more precise control is available. Existing methods are difficult to precisely control because particles can be located at various positions in the pressure field other than the target position. However, by using the chirp waveform, it is possible to control only the particles in the desired specific position.
Using a waveform with time-varying frequency instead of using a fixed-frequency continuous waveform, provides distinct advantages over current methods for neuromodulation and ARF. Chirp is one example to improve the spatial resolution. Another variation is the location of the transducer. In this description, two transducers are facing to each other. However, two or more transducers located at different angular locations can be used.
In prior ultrasound neuromodulation methods, stimulation was performed in all areas where high pressure was applied. As reducing the focal spot size was the only way to increase the spatial resolution for ultrasound neuromodulation, all the studies concentrated in reducing the focal spot size of the focused transducer. In contrast, the present invention teaches that, in the case of PIRF with chirp waveform, it is possible to reduce the stimulating area without reducing the size of the focal spot. This method can increase the spatial resolution of neuromodulation without reducing the size of the focal spot. This technique also allows the position of the neuromodulation spot to be moved electronically by giving a time delay to a single transducer. This allows moving the neuromodulation spot without moving the focal points of the fixed transducers.
This application is a continuation-in-part of U.S. patent application Ser. No. 16/894,421 filed Jun. 5, 2020, which claims priority from U.S. Provisional Patent Application 62/858,884 filed Jun. 7, 2019, which are both incorporated herein by reference. This application also claims priority from U.S. Provisional Patent Application 63/326,669 filed Apr. 1, 2022, which is incorporated herein by reference.
This invention was made with Government support under contract N5112152 awarded by the National Institutes of Health, and under contract EB023901 awarded by the National Institutes of Health. The Government has certain rights in the invention.
| Number | Date | Country | |
|---|---|---|---|
| 62858884 | Jun 2019 | US | |
| 63326669 | Apr 2022 | US |
| Number | Date | Country | |
|---|---|---|---|
| Parent | 16894421 | Jun 2020 | US |
| Child | 18129680 | US |
