Acute Damage Threshold for Infrared Neural Stimulation of the Cochlea: Functional and Histological Evaluation
Article first published online: 8 OCT 2012
Copyright © 2012 Wiley Periodicals, Inc.
The Anatomical Record
Special Issue: The Anatomy and Biology of Hearing and Balance: Cochlear and Vestibular Implants
Volume 295, Issue 11, pages 1987–1999, November 2012
How to Cite
Goyal, V., Rajguru, S., Matic, A. I., Stock, S. R. and Richter, C.-P. (2012), Acute Damage Threshold for Infrared Neural Stimulation of the Cochlea: Functional and Histological Evaluation. Anat Rec, 295: 1987–1999. doi: 10.1002/ar.22583
- Issue published online: 22 OCT 2012
- Article first published online: 8 OCT 2012
- Manuscript Received: 24 JUL 2012
- Manuscript Accepted: 24 JUL 2012
- The National Institute on Deafness and Other Communication Disorders
- National Institutes of Health
- Department of Health and Human Services. Grant Number: HHSN260-2006-00006-C/NIH No. N01-DC-6-0006
This article provides a mini review of the current state of infrared neural stimulation (INS), and new experimental results concerning INS damage thresholds. INS promises to be an attractive alternative for neural interfaces. With this method, one can attain spatially selective neural stimulation that is not possible with electrical stimulation. INS is based on the delivery of short laser pulses that result in a transient temperature increase in the tissue and depolarize the neurons. At a high stimulation rate and/or high pulse energy, the method bears the risk of thermal damage to the tissue from the instantaneous temperature increase or from potential accumulation of thermal energy. With the present study, we determined the injury thresholds in guinea pig cochleae for acute INS using functional measurements (compound action potentials) and histological evaluation. The selected laser parameters for INS were the wavelength (λ = 1,869 nm), the pulse duration (100 μs), the pulse repetition rate (250 Hz), and the radiant energy (0–127 μJ/pulse). For up to 5 hr of continuous irradiation at 250 Hz and at radiant energies up to 25 μJ/pulse, we did not observe any functional or histological damage in the cochlea. Functional loss was observed for energies above 25 μJ/pulse and the probability of injury to the target tissue resulting in functional loss increased with increasing radiant energy. Corresponding cochlear histology from control animals and animals exposed to 98 or 127 μJ/pulse at 250 Hz pulse repetition rate did not show a loss of spiral ganglion cells, hair cells, or other soft tissue structures of the organ of Corti. Light microscopy did not reveal any structural changes in the soft tissue either. Additionally, microcomputed tomography was used to visualize the placement of the optical fiber within the cochlea. Anat Rec, 2012. © 2012 Wiley Periodicasl, Inc.
Infrared neural stimulation (INS) has been proposed as a novel method to directly stimulate neurons, their axons, or their dendrites (Wells et al., 2005a, b; Izzo et al., 2006b). Pulses of mid-infrared (IR) laser radiation are delivered to native, nonmodified neural tissue and result in neural depolarization. The main advantage of INS is the possibility of stimulating neuron populations with a high spatial resolution when compared with electrical stimulation (Teudt et al., 2007; Izzo et al., 2007b; Wells et al., 2007b; Matic et al., 2011; Richter et al., 2011b). At the wavelengths used for INS, the energy is primarily absorbed by water in the tissue and is converted into heat (Wells et al., 2006; Wells et al., 2007a). Recently, it has been demonstrated that local heating of the neural tissue by the laser pulse induces a capacitive and depolarizing current at the cell membrane (Shapiro et al., 2012). Although INS has its advantages, each laser pulse causes a transient temperature increase in the surrounding tissue and accumulation of the thermal energy could result in injury to the target or surrounding tissue. At pulse repetition rates that would allow acoustical information to be encoded in a cochlear implant, it is important to determine the safe energy range over which cochlear function is not adversely affected. Deposition of thermal energy could result in injury to the cochlear structures and could reduce efficacy of the implant. Previous experiments that explored short-term sustainability (<10 hr of continual stimulation) of optical stimulation were conducted in the gerbil at 10 Hz repetition rate (Izzo et al., 2006b) and in the cat models (Rajguru et al., 2010). The experiments have shown that IR irradiation with the selected laser parameters did not result in thermal damage of the cochlea. Experiments during which cats were chronically implanted and cochlear function was monitored over at least 6 weeks did not reveal any changes in cochlear function either (Robinson et al., 2011). The previous studies did not attempt to systematically evaluate the laser parameters that will result in thermal injury.
In addition to a short review on the recent advances in the field of INS, this manuscript provides results from experiments that determined the acute damage threshold of the cochlea for INS. The cochlea was continuously irradiated with IR pulses, which were delivered by a single optical fiber that was coupled to a laser. For increasing radiant energies and a fixed pulse repetition rate that allows the encoding of acoustical information in a cochlear implant, cochlear function was monitored over time using the compound action potential (CAP) amplitude. Cochlear electrophysiological recordings are a potent marker for the state of the cochlea, often showing changes in amplitude or morphology to very minor perturbations or damage. In addition, histology was performed to examine possible thermal damage to the tissue.
Adult guinea pigs (300–600 g) of either sex were used for the experiments. All procedures were carried out in accordance with the NIH Guide for the Care and Use of Laboratory Animals and were approved by the Animal Care and Use Committee of Northwestern University. The procedures have been described in detail previously (Moreno et al., 2011; Richter et al., 2011b).
Briefly, the cochlea was surgically accessed and a cochleostomy was created. An optical fiber coupled to an IR laser was advanced through the cochleostomy into the scala tympani. An electrode was placed at the round window to record the cochlear CAP from the cochlea evoked in response to optical pulses from the laser. Before the cochlea was stimulated continuously with INS, a sequence of INS-evoked CAPs was recorded to determine baseline cochlear function. Thereafter, the cochlea was stimulated continuously and cochlear function was monitored with optically evoked CAPs. A drop in CAP amplitude by 25% or a total recording time of ≥3 hr was the criterion for “damage” and “no damage,” respectively. The cochleae were stimulated at different radiant energies. At the conclusion of the experiments, the cochleae were harvested for histology.
Animal Anesthesia and Surgery
Animals were anesthetized with urethane (1.3 g/kg i.p. in 20% sterile phosphate-buffered solution [PBS]). Urethane injections were supplemented with ketamine (44 mg/kg) and xylazine (5 mg/kg) at the beginning of the surgical procedures. Depth of anesthesia was assessed every 15 min with a paw withdrawal reflex. Anesthesia in all guinea pigs was maintained by supplements of ketamine (44 mg/kg) and xylazine (5 mg/kg) along with saline solution (0.5 mL). To reduce bronchial secretions, atropine sulfate (0.04 mg/kg) was administered at the beginning of the experiment. Core body temperature was maintained at 38°C with a thermostatically controlled heating pad. After the animals were anesthetized, a tracheotomy was made and a plastic tube (1.9 mm outer diameter, 1.1 mm inner diameter, Zeus, Orangeburg, SC) was secured into the trachea. The tube was connected to an anesthesia system (Hallowell EMC, Pittsfield, MA). The animals were ventilated on oxygen throughout the length of the experiment.
After the animals were anesthetized, a c-shaped skin incision was made behind the left ear lobe and the cervicoauricular muscles were removed. The cartilaginous outer ear canal was exposed and cut to insert an ear bar into the ear canal. A hollow ear bar on the left side and a solid ear bar on the right side were used to fix the head in a stereotactic head holder (Stoelting, Kiel, WI). The hollow bar allowed acoustic stimulation of the left ear, and the solid ear bar blocked the right ear canal, in addition to maintaining the head in a fixed position. The left bulla was exposed and opened approximately 2 × 3 mm with a motorized drill (World Precision Instruments, Sarasota, FL). The basal turn of the cochlea was identified and a cochleostomy was created with a 0.5-mm Buckingham footplate hand drill (Richards Manufacturing, Memphis, TN). An optical fiber (P200-5-VIS-NIR, Ocean Optics, Dunedin, FL) was inserted through the opening of the cochlear wall. For the present experiments, the fiber was 200 μm in core diameter, the numerical aperture was 0.22 and the acceptance angle was 25.4° in air and about 2° in fluids. The individual optical fibers were mounted to a micromanipulator (MHW103, Narishige, Tokyo, Japan) to ensure consistent orientation during stimulation. To measure INS-evoked CAPs, a silver ball electrode was placed on the round window and a reference electrode was placed under the skin at the neck.
Stimulus Generation and Calibration
Voltage commands for acoustic pure tone stimuli were generated using a computer I/O board (KPCI 3110, Keithley, Cleveland, OH) inserted into a PC and were used to drive a Beyer DT 770Pro headphone (Beyerdynamic, Farmingdale, NY). The speaker was coupled with a short, 3-mm diameter plastic tube to the opening of the hollow ear bar. Acoustic stimuli were tone pips (12 ms duration, including a 1 ms rise/fall) with different carrier frequencies, which were presented at a rate of 4 Hz. The resulting sound level at the end of the hollow ear bar was measured with a 1/8-in microphone (Bruel & Kjaer North America, Norcross, GA). For the measurements, the microphone's protective grid was flush at the opening of the ear bar.
IR Neural Stimulation
INS was achieved with a diode laser (Capella, Lockheed Martin Aculight, Bothell, WA). For the present experiments, the wavelength was 1,869 nm and the pulse duration was 100 μs. The laser was operated at 200–500 Hz repetition rate and was coupled to the optical fiber. The spot size at the tip of the optical fiber was 130 μm in diameter (full width at half maximum, FWHM), with a Gaussian energy distribution. The FWHM value was determined previously using the knife edge technique as described previously (Teudt et al., 2007). The energy per pulse at the tip of the optical fiber was measured in air with the J50LP-1A energy sensor (Coherent, Santa Clara, CA) and was 0–127 μJ/pulse. However, the energy measured at the tip of the optical fiber will not be the same as the energy delivered to the neurons. Fluids, soft tissue, and the modiolar bone absorb and scatter the radiation. The penetration depth of the radiation at 1,869 nm is about 600 μm, assuming primarily water absorption (e.g., Hale and Querry, 1973). Using postexperimental reconstructions of the stimulated cochleae with the microcomputed tomography (microCT) scans, distances between the optical fiber and the spiral ganglion cells were determined. As the incident energy decreases in water by 1/e (63%) for each 600 μm traveled along the optical path, it is a fair assumption that the energy at the spiral ganglion cells is about one-third of the energy measured at the tip of the optical fiber. Energy values provided in this article are the radiant exposures that were measured at the tip of the optical fiber in air.
Data Acquisition and Analysis
INS-evoked CAPs were recorded with the round window electrode, which was connected to a differential amplifier (ISO-80, WPI, Sarasota, FL). The input-impedance was >1012 Ω and the gain was set to 60 dB. The responses were bandpass filtered (0.3–3 kHz). The sampling rate was 25 kHz, and a minimum of 10 consecutive pulse-by-pulse responses to INS at were averaged for each data point. CAP peak-to-peak (P1–N1) amplitude was monitored and graphed versus time. A CAP baseline was established at a fixed energy between 25 and 127 μJ/pulse. For some animals, the responses to stimulation at 10 Hz repetition rate were monitored in addition to the responses obtained at 250 Hz pulse repetition rate. CAP amplitudes were subsequently normalized to the baseline.
With high-resolution microCT, objects with at least 50–100 μm spatial resolution can be obtained. MicroCT has been used previously to image the cochlea (Vogel, 1999; Rau et al., 2006; Lareida et al., 2009; Richter et al., 2009; Glueckert et al., 2011; Smith et al., 2011). This imaging method uses individual radiographs or projections that were captured from different viewing directions to reconstruct the internal structure of the object. In the present experiments, the method was used to determine the internal bony structures of the guinea pig cochlea and the location of the optical fiber that was inserted through a cochleostomy into the cochlea. The method is noninvasive. A detailed description of the Methods is provided by Stock (2009).
Prior to scanning the cochleae using the MicroCT-40 system (Scanco Medical, Wayne, PA), the optical fiber was cemented in place after the final physiological measurement. Approximately, 1 mL of dental acrylic was prepared in a small glass container and was aspirated with a 1-mL syringe. A 20-gauge needle was attached to the syringe and the dental acrylic was then carefully injected into the bulla of the guinea pig. Care was taken not to move the optical fiber during the procedure. After the acrylic had cured, the animals were euthanized and the bulla were removed from the skull and were placed into paraformaldehyde (4% RL) for a minimum of 24 hr. The cochleae were transferred and washed in RL and subsequently were placed in a plastic tube for scanning. The cochleae were imaged with the microCT operated at 45 kV tube voltage, 88 μA tube current, and 300 ms integration time for each projection. A complete description of the scanner can be found elsewhere (Scanco Medical Web site, www.scanco.ch). Each of the 120–300 contiguous slices was reconstructed on 1,024 × 1,024 grid with 30 μm isotropic volume elements (voxels). The reconstructed slices were imported into ImageJ (http://rsbweb.nih.gov/ij/) and were converted into a stack of TIFF files. The images were segmented using the Amira Software (Visage Imaging, San Diego, CA). A representation in three dimensions was created for scala tympani, spiral ganglion, cochlear nerve, and optical fiber.
In addition to using CAPs as the functional marker for injury thresholds, we evaluated histological sections following INS. Animals were euthanized and the cochleae were harvested and subsequently fixed with paraformaldehyde (4% in PBS solution). To decalcify the cochleae, they were placed in 10% ethylenediaminetetraacetic acid for 7–14 days. Following decalcification, the cochleae were rinsed in PBS three times for 15 min each and dehydrated in a graded six-step acetone series (1 hr each in 25, 50, 75, 90, 100, and 100% acetone). Cochleae were embedded in Araldite-Epoxy Resin (7:1 acetone:resin, 1:1 acetone:resin, 1:7 acetone:resin, one-time pure Resin). The plastic was cured for 12 h in an oven at 60°C.
The cochleae were cut perpendicular to the modiolus at 5 μm thickness using a ultramicrotome (LKB 8800 Ultrotome III, Stockholm-Bromma, Sweden) and the sections were placed on glass slides and were stained with 1% toluidine blue (Sigma Aldrich, St. Louis, MO) and 1% sodium tetraborate solution (1:20). Images of each series of consecutive sections were captured with a microscope (Zeiss) equipped with a Pixar 1000 color camera. The sections were used to identify signs of thermal damage. The following parameters were investigated: visible damage, readily apparent lesion with some circular symmetry, severe lesion, and circular symmetric opacity with shrinking of epithelium at the center. Furthermore, the presence of inner and outer hair cells, inner and outer pillar cells, supporting cells, and the tectorial membrane and basilar membrane was documented for each section. Spiral ganglion cells along corresponding segments of the basal turn of the cochlea were counted using a stereological method as described previously (Richter et al., 2011a).
Averages and standard deviations were calculated for the CAP amplitudes and for the cell numbers obtained from the histological images. ANOVA was performed, followed by the Tukey test, to determine whether differences observed were significant.
Placement of the Optical Fiber
In eight specimens, the placement of the optical fiber was determined. As described in METHODS section, the optical fiber was cemented in place after the physiological measurements were completed. The bulla were removed and imaged in a microCT scanner. Figure 1A shows the magnified view of the basal turn of the guinea pig cochlea. The different key structures in Fig. 1A have been identified in a sketch shown in Fig. 1B. The slice of the cochlea (Fig. 1C) shows the cochleostomy and the optical fiber. Orientation and distance for a typical optical fiber placement relative to the neural structures were determined from this section and were between 300 and 1,067 μm with an average of 857 μm (N = 8). The inset in Fig. 1C shows the optical fiber in a different slice. MicroCT data were imported into the Amira Software (Visage Imaging, San Diego, CA) and key structures of the cochlea were rendered in 3D (Fig. 1D), including scala tympani (transparent purple), the spiral ganglion (red), the auditory nerve in the modiolus (green), and the optical fiber (yellow).
Cochlear damage was defined by a decrease in CAP amplitude of more than 25% when compared to the baseline. Figure 2 shows an example for continuous stimulation at a radiant energy of 127 μJ/pulse and at repetition rates of 10 and 250 Hz. Raw (Fig. 2A) and normalized (Fig. 2B) data are shown. After a baseline was established at the beginning of the experiment, the cochlea was stimulated continuously for 3 hr at 250 Hz pulse repetition rate. For this example, no significant change in CAP amplitude occurred during ∼3 hr of continuous stimulation. To test whether cochlear damage can be achieved at repetition rates above 250 Hz, in this single case, the pulse repetition rate was increased to 500 Hz. At the higher rate, the CAP amplitude decreased within 1 hr so that it was indistinguishable from the noise floor of the measurement (data not shown).
In N = 29 cochleae, the radiant energy per pulse was varied systematically, whereas the pulse repetition rate was fixed at 250 Hz. The cochleae were stimulated for at least 3 hr. For cochleae, in which the CAP amplitude decreased by 25% within 3 hr, the time between the start of stimulation and the time of 25% reduction has been graphed in Fig. 3. We did not observe any change in the CAP amplitude within 3 hr for continuous INS when the radiant energy was below 30 μJ/pulse. For radiant energies of 30 μJ/pulse or larger, the times for a 25% reduction in CAP amplitude decreased with increasing radiant energy (Fig. 3). The data were used to calculate the likelihood of cochlear damage by dividing the number of cochleae that showed damage at a given radiant energy by the number of cochleae tested at the particular energy (Fig. 4). Hence, 1 indicates that all cochleae exposed to the selected energy were damaged and 0 indicates that none of the cochleae was damaged. From the data, we observed that there was no likelihood for cochlear damage at 25 μJ/pulse, but it drastically increased if the radiant energy reached 40 μJ/pulse (Fig. 4).
Three cochleae that were not exposed to INS were selected randomly as controls along with five cochleae that were stimulated with IR laser pulses. The laser pulses were 100 μs in duration, the radiant energy was 127 μJ/pulse (N = 3) and 98 μJ/pulse (N = 2), and the repetition rate was 250 Hz. All exposed cochleae selected for histology showed a decrease in CAP amplitudes and were classified as damaged cochleae. From the selected cochleae, 15–20 consecutive sections of the mid modiolar plane, targeted by the beam path, were used for the analysis. Each of the sections was inspected visually. The light microscopic appearance of the organ of Corti and the spiral ganglion neuron count in the exposed areas along the cochlea did not change. Outer and inner hair cells, the tectorial membrane and the basilar membrane could be identified in all sections obtained from the cochleae. The cochlear structures were similar in the irradiated and the control cochleae. The sections did not show structural alterations of the tissue, including edema, cell shrinkage, complete fragmentation of the nucleus into discrete bodies, vacuole formation, or signs of coagulation. Spiral ganglion neuron counts did not differ significantly between the groups. Note that the six exposed cochleae were grouped together.
For the interpretation of the histological images, it is important to remember that the survival time after the exposure was <1 hr. Therefore, changes that are caused by apoptosis and necrotic degeneration are not expected to be observable with light microscopy in the selected sections.
INS—A Review of the Current Advances in the Field
One goal of research using artificial neural stimulation is to design devices that are able to restore or replace neural function. Contemporary neural prostheses restore motor function in patients who have suffered from stroke, cochlear implants restore hearing in severe-to-profound deaf individuals, vestibular prostheses are designed to restore balance and treat vertigo, and deep brain stimulators aim to decrease the symptoms of Parkinson's disease. The devices stimulate the neurons with electrical current, which spreads in tissue and renders spatially selective stimulation difficult. Electrical stimuli also produce stimulation artifacts making stimulation and simultaneous recording of a neural response challenging. Furthermore, the electrodes often have direct contact with the tissue, which can result in tissue damage caused by the physical contact of the electrode or by large current densities generated close to the electrode.
It has recently been suggested that many of these limitations may be overcome by INS using pulsed near-IR radiation from a laser. The idea of exciting cells using light dates back to 1891 (Arsonval, 1891) and the first reports of photostimulation of neural tissue were published as early as 1947. Arvanitaki and Chalazonitis (1961) demonstrated that visible light at different wavelengths could produce effects in the visceral ganglia of Aplysia. However, these experiments were conducted within a more limited wavelength and temporal range than is available from optical sources today.
A novel optic-based method of neural stimulation has been explored over the last 10 years. Tissue is irradiated with pulses of near-IR radiation, which depolarize neurons in the optical path. This method of neural stimulation, termed INS, does not require the introduction of a chromophore or light-gated ion channels, as is necessary for optogenetics (for a review, see Kramer et al., 2009). It does not need contact between the optical source and the neural tissue. Most of the radiation wavelengths used for these studies (λ = 1,470–1,550 nm; λ = 1,840–2,120 nm) can be delivered via an optical fiber, which provides flexibility of minimally invasive delivery. Furthermore, simultaneous optical stimulation and electrical recording is possible, as there is no stimulation artifact present in the data. However, the most attractive feature of INS is the improvement in spatial selectivity of stimulation. The radiation does not spread significantly in the tissue, compared to electric current, and can be further focused using lenses. The ability to confine the optical stimulus to a smaller portion of neurons may allow for more discrete neural stimulation with improved resolution. Although there are many research questions still to be answered regarding INS, the initial results are promising. Below, we present a summary to-date on pulsed, mid-IR stimulation of neurons.
Stimulation of Peripheral Nerves
Wells et al. (Wells et al., 2005a, b) used a pulsed IR laser to stimulate the sciatic nerve and have pioneered a new field of artificial neural stimulation using optical radiation. They successfully evoked action potentials from the rat sciatic nerve in response to laser pulses without the use of exogenous fluorophores or genetically modified neurons. The stimulation sources for their studies were a free electron laser (λ = 2,000–6,000 nm; τp = 5 μs), a Ho:YAG laser (λ = 2,120 nm; τp = 350 μs) at low energies, and pulsed diode lasers (λ = 1,844–1,937 nm; τp = 10 μs to 10 ms; Lockheed Martin Aculight). It was demonstrated that neural stimulation is possible at optical wavelengths that are moderately absorbed in neural tissue. The estimated optical penetration depth in neural tissue is 200–1,500 μm (Falk and Ford, 1966; Hale and Querry, 1973), assuming primarily water absorption (Rosenberg and Gunner, 1959; LoPachin and Stys, 1995). It has also been demonstrated that wavelengths with similar penetration depths (1,540, 1,495, and 1,450 nm) can be used for INS (McCaughey et al., 2009). However, the pulse durations required to depolarize the cells at these wavelengths were significantly longer (∼100 ms) and the fluences required for stimulation were higher than those demonstrated with 1,844–1,937 nm. This may limit the potential applications of 1,450–1,540 nm radiation in neuroprostheses. Researchers have recently investigated a hybrid electrical and INS system, to determine if there are any inhibitory or synergistic effects of the two stimulation modes. When using a subthreshold electrical pulse, the optical energy required to evoke a compound muscle action potential via INS decreased, as compared to INS alone (Duke et al., 2009).
INS has also been demonstrated on the rat nervus pudendus. A thulium fiber laser (λ = 1,870 nm; τp = 2.5 ms; f = 10 Hz) was used to stimulate the nerves with 7.5 mJ/pulse. During the experiments, the nerves were stimulated for 60 sec blocks of time (Fried et al., 2008a, b). A significant increase in the intracavernosal pressure was detected during INS. Although a laser is unlikely to be used to restore male sexual function, INS may become an important method used to identify and preserve the pudendal nerve during prostate surgery.
It has been hypothesized that the mechanism of stimulation is a local, transient temperature increase from light absorption by water in the tissue (Wells et al., 2007a). A temperature accumulation, however, may lead to injury and limit the efficacy of INS. Wells et al. (Wells et al., 2007c) have previously explored whether INS of the sciatic nerve results in nerve damage. The upper limit for safe laser stimulation in peripheral nerves occurred near 5 Hz pulse repetition rate and the maximum duration for constant low repetition rate stimulation (2 Hz) was approximately 4 min with adequate tissue hydration. This parameter space would render the method unusable in neural interfaces that require higher repetition rates and long-term stimulation, such as with cochlear implants. However, in contrast to stimulation of the peripheral nerves, short- and long-term experiments in the auditory system showed that stimulation rates of 200 Hz over 8 hr to 6 weeks did not result in noticeable functional changes (see below for details).
Stimulation of Cranial Nerves
INS has also been used to stimulate the facial nerve (Teudt et al., 2007), the auditory nerve (Izzo et al., 2006a; Izzo et al., 2007b, d; Izzo et al., 2008), and the vestibular crista ampullaris (Rajguru et al., 2011). The facial nerve was irradiated using a 600 μm-diameter optical fiber with radiant exposures between 0.71 and 1.77 J/cm2 and the evoked compound muscle action potentials were measured from the facial muscles (Teudt et al., 2007). The stimulation was selective and did not result in observable neural damage with radiant exposures up to 2 J/cm2 over the short term.
Optical stimulation of the auditory nerve has been studied in normal hearing gerbils (Izzo et al., 2006a, b; Izzo et al., 2007a–d; Izzo et al., 2008; Matic et al., 2011), deafened gerbils (Richter et al., 2008), the mouse (Suh et al., 2007), the guinea pig (Moreno et al., 2011; Richter et al., 2011b), and the cat (Rajguru et al., 2011). In a first series of experiments, a Ho:YAG laser was used to stimulate the auditory nerve of gerbils (Izzo et al., 2006b). For neural stimulation, the laser was coupled to a 200 μm-diameter optical fiber, which was placed in front of the cochlear round window in gerbils. The optical beam was directed toward the spiral ganglion cells and compound action potentials were evoked directly in response to the cochlear INS. An increase in laser energy induced a monotonic increase in the evoked response. Subsequent experiments with different diode lasers (λ = 1,844–1,937 nm; τp = 5 μs–1 ms; f = 2–1,000 Hz) showed that the patterns of the compound action potentials changed with increasing pulse duration (Izzo et al., 2007d; Izzo et al., 2008). In acutely and chronically deaf gerbils relative to normal hearing controls, there was little change in energy thresholds at short pulse durations (Richter et al., 2008). Long-term acute stimulation was achieved for 6 hr in the gerbils at 13 Hz (Izzo et al., 2007d) and 8 hr in cats at 200 Hz (Rajguru et al., 2010) and showed stable CAP amplitudes for the duration of stimulation.
Experiments are currently underway to investigate the safety and efficacy of a chronic optical cochlear implant in a cat model. We have developed a method to chronically implant a multichannel light delivery system that can be used to stimulate the animals over extended periods of time. These data will verify that the smaller spread of excitation can result in a separation of the independent channels to stimulate the cochlea. The experiments will also provide safe repetition rates and energy levels for chronic stimulation. It remains to be seen if the neural responses evoked by the optical stimulus correlate with a meaningful perception of sound. Preliminary results obtained with a single channel implant do not suggest any damage caused by the long-term stimulation (Robinson et al., 2011). The implant was in place for several weeks and was used to stimulate the cochlea for 6 hr per day during this time. It should be emphasized that that the design of the light delivery system as presented is not a prototype for a human cochlear implant based on INS. The current design serves to determine the basic parameters for laser stimulation that can guide the development of a future implantable human prototype. INS for the cochlea can take advantage of the tonotopic organization of the cochlea and stimulate selected neuron populations along the cochlea to convey frequency information. Crossturn stimulation, as has been shown in the gerbil cochlea, likely will not occur in a larger cochlea, such as the cat or the human cochlea. The distance across turns is in the millimeter range and the fluids and tissue structures will absorb the radiation.
The selectivity of INS in the auditory system has been investigated via several methods in gerbils and guinea pigs. The results suggest that a single channel optical stimulus is more selective than a monopolar electrical stimulus and produces spatial tuning curves (STCs) on the order of selectivity of a low-level acoustic tone (Izzo et al., 2007b; Matic et al., 2011; Richter et al., 2011b).
Another potential application of INS has been suggested in vestibular prostheses, as a means to reduce imbalance and disorientation caused by vestibular dysfunction. Harris and coworkers (Harris et al., 2009) showed that optical stimulation with 1,840 nm radiation (τp = 10 μs–1 ms) was indeed effective at evoking compound nerve potentials when stimulating at the VIIIth nerve. At the simulation site selected, the nerve is comprised of both auditory and vestibular afferents. However, optical stimulation of the ampullae did not evoke detectable eye movements. In a different set of experiments, Scarpa's ganglion in gerbils was stimulated with diode lasers (λ = 1,844–1,877 nm; τp = 5 μs–1 ms; f = 2–1,000 Hz), while recording neural activity from single vestibular nerve fibers (Bradley, 2009). The experiments showed an evoked compound action potential but did not result in detectable changes of single fiber activity.
Recently, it has been shown that INS can be used to release neurotransmitter from vestibular hair cells and to modulate the discharge rate of primary afferent neurons in the VIIIth cranial nerve (Rajguru et al., 2011). The lateral semicircular canal ampullary organ of the oyster toadfish was used as the experimental model. Single-unit postsynaptic afferent recordings during pulsed irradiation of hair cells revealed three characteristic response types: excitatory “on” units that dramatically increased discharge rate, inhibitory “off” units that decreased discharge rate or were completely silenced, and mixed units that initially decreased discharge rate and subsequently increased discharge rate. The phasic responses developed within 8 ms, whereas tonic afferent responses developed with a time constant of ∼7 ms–10 s following the onset of the INS at the hair cell. Recovery to the prestimulus discharge rate followed a similar time course. In addition to the relatively slow changes in discharge rate, a subset of afferent neurons responded with an action potential for each INS pulse and phase-locked their discharge to the stimulation rate. The results demonstrate that phase locking occurred for 10–80 Hz INS with a latency around 7.6 ms (Rajguru et al., 2011). Although the data supported the hypothesis that IR evoked neurotransmitter release resulted in afferent responses recorded, evoked eye movements were not measured. It remains to be seen if pulsed INS of the vestibular end organs can indeed generate physiological eye movements and find an application in vestibular prostheses.
Stimulation of the Brainstem
Lee and coworkers (Lee et al., 2009) have demonstrated that neurons of the central auditory system can be stimulated via INS. In acute animal experiments, a 400 μm-diameter optical fiber was placed on the surface of the cochlear nucleus to irradiate the tissue with diode lasers (λ = 1,849–1,865 nm; τp = 5 μs–10 ms; f = 2–1,000 Hz). An optically evoked auditory brainstem response (oABR) was recorded and resembled a multipeaked ABR evoked by acoustic stimulation. The oABR waveforms had latencies between 3 and 8 ms, longer than ABRs evoked by direct electrical stimulation of the same region. Reproducible oABRs were evoked at thresholds as low as 169 mJ/cm2, with 50 μs pulse duration and 5 Hz repetition rate. The oABR was stable during continuous INS for 30 min. Control experiments, in which the optical path was blocked and the rat was euthanized, eliminated the oABR. No thermal tissue damage was found on histological examination when stimulating with pulse durations <1 ms and radiant exposures <2.05 J/cm2. While promising, further experiments are required to show that INS is effective at higher repetition rates and that there is improved spatial selectivity at the cochlear nucleus.
Stimulation of the Cortex
Recently, the cortex has been successfully stimulated using INS, using an in vitro preparation of thalamocortical brain slices (Cayce et al., 2010) and in vivo stimulation of the somatosensory cortex (Cayce et al., 2011). In the latter study, pulsed IR light (λ = 1,875 nm, pulse duration = 250 μs, optical fiber core diameter = 400 μm, repetition rate = 50–200 Hz) was used to stimulate the somatosensory cortex of anesthetized rats and its efficacy was assessed using intrinsic optical imaging and electrophysiology techniques. INS was found to evoke an intrinsic response of similar magnitude to that evoked by tactile stimulation. A maximum deflection in the intrinsic signal was measured to range from 0.05 to 0.4% in response to INS, and the activated region of cortex measured approximately 2 mm in diameter. The magnitude of the intrinsic signal increased with higher laser repetition rates and increasing radiant exposures. Single-unit recordings indicated a significant decrease in neuronal firing that was observed at the onset of INS stimulation and lasted for up to 1 sec after stimulation onset. The pattern of neuronal firing differed from that observed during tactile stimulation, potentially owing to a different spatial integration field of the pulsed IR light compared to tactile stimulation.
Stimulation of Other Excitable Cells
In addition to neurons or nerves, it has been demonstrated that other excitable cells can also be stimulated with IR. It has been shown that femtosecond pulses of focused near-IR laser radiation caused contractions in cultured neonatal rat cardiomyocytes (Smith et al., 2008). By periodic exposure to pulse trains, contraction cycles in cardiomyocytes could be triggered and synchronized with the laser. This was observed in isolated cells and in small groups of cardiomyocytes with the laser acting as pacemaker for the entire group. The range of average radiant power to achieve this effect was between 15 and 30 mW for an 80 fs, 82 MHz pulse train at 780 nm. At higher power levels, cells typically demonstrated a large intracellular calcium elevation and contracted without subsequent relaxation. This laser–cell interaction allowed the laser radiation to act as a pacemaker, and could be used to trigger contraction in dormant cells as well as synchronize or destabilize contraction in spontaneously contracting cardiomyocytes (Smith et al., 2008). In a more recent study, responses of neonatal rat ventricular cardiomyocytes to pulsed IR radiation (λ = ∼1,862 nm) were studied. Fluorescence monitoring of the intracellular free calcium (Ca2+) demonstrated that IR radiation induced rapid (millisecond time scale) intracellular Ca2+ transients in the cells (Dittami et al., 2011). The results showed that the Ca2+ transients were sufficient to elicit contractile responses from the cardiomyocytes and the cardiomyocytes could be “paced” or entrained to the laser stimulation rate. Pharmacological results implicate mitochondria as the primary intracellular organelles contributing to the laser-evoked Ca2+ cycling (Dittami et al., 2011). Light has also been used to optically pace an intact heart in vivo (Jenkins et al., 2010). Pulsed 1,875-nm IR laser light was used to pace the heart rate to the stimulation rate of the laser. Laser doppler velocimetry was used to verify the pacing. At low radiant exposures, embryonic quail hearts were reliably paced in vivo without detectable damage to the tissue, indicating that optical pacing has great potential as a tool with which to study embryonic cardiac dynamics and development. In particular, optical pacing can be used to control the heart rate, thereby altering stresses and mechanically transduced signaling.
Advantages of Optical Stimulation
Optical manipulation of excitable cells via endogenous sensitivity or transfection of proteins is an important tool today in neuroscience. The use of lasers to stimulate neurons and other excitable cells is of increasing interest, especially given the success of pulsed and continuous lasers in depolarizing neurons. The approach used here is pulsed IR radiation to control cell excitability via endogeneous mechanisms, which does not require chemical manipulation of the target cells and it may also have several advantages over traditional electrical stimulation. Optical stimulation yields artifact-free signals, which enables the simultaneous recording of neural responses and would allow for long-term evaluation of the state of the tissue. This is difficult to achieve with electrical stimulation. INS can also be delivered in a noncontact manner. The observations from gerbils and cats, where long duration stimulation did not affect the amplitude of evoked compound action potentials, lend further support to the safety of lasers in chronic implant use.
The main advantage of INS is its spatial selectivity, when compared with conventional electrical stimuli. In the sciatic and facial nerve studies, different nerve bundles within the main nerve trunk were individually stimulated when the radial location of the stimulation site was varied (Teudt et al., 2007; Wells et al., 2007b). This is in contrast to electric stimulation, which evoked strong and unselective responses in all muscle groups innervated by the nerve. In the inner ear, the transiently expressed transcription factor, c-FOS, was used to stain activated nerve cells to identify the spatial area of the cochlea that was stimulated. The results showed that INS of the cochlea was more selective compared to electrical stimulation. At high stimulus levels, stained cells were in the beam path determined, which was estimated from the placement of the optical fiber (Izzo et al., 2007b). Electrical stimulation at high levels, however, resulted in the activation of all spiral ganglion cells in the cochlea (Izzo et al., 2007b). Furthermore, tone-on-light masking experiments and recordings from the inferior colliculus (ICC) demonstrated that the laser could stimulate a small population of cells similar to an acoustic tone burst (Matic et al., 2011; Richter et al., 2011b). INS-evoked STCs that were constructed from the neural activity recorded from central nucleus of the ICC showed single peak and multilobe responses. For approximately 13% of the optically evoked ICC STCs, multiple peaks were seen. As has been discussed by Richter et al. (2011), it may be argued that multiple peaks seen on the STCs are caused by structures in the optical path that block the radiation from stimulating the neurons. In a separate series of experiments, reconstructions of the optical path have been made from serial sections of the guinea pig cochlea (Moreno et al., 2011). The optical paths obtained from the reconstructions were then compared with the ICC neural responses obtained in the same animal. Based on the results of the reconstructions, it is not clear whether bony structures in the optical path blocked the radiation and produced “shadows” in stimulation. Examining the ICC neural activity suggested that these profiles, for most of the cases, actually represented one broad response area and that the “gaps” seen in the STCs are an artifact of the data analysis method. In contrast to the multipeak responses, approximately 55% of the STCs evoked by INS of the cochlea were narrow. The orientation of the optical fiber toward the spiral ganglion and the resulting optical path has a direct impact on the location of the ganglion being stimulated and the spread of stimulation. For all the experiments, the access to the basal turn of the cochlea was at a similar location. However, the site of stimulation varied along the cochlea as the orientation of the optical fiber was not held constant. Stimulation sites between 8 and 16 kHz lay directly in front of the optical fiber inserted through the basal turn cochleostomy and the optical path for these sites was perpendicular to the spiral ganglion. To target lower or higher frequencies, the optical fiber had to be inserted at an angle into the scala tympani such that the stimulation site was within the optical path. In that case, the orientation of optical path would be more tangential to the spiral ganglion, likely resulting in a larger segment of the spiral ganglion being irradiated. A broader STC would, therefore, be expected (Richter et al., 2011b).
Single-peak optically evoked STCs also showed best frequencies that do not correlate with locations directly in front of the optical fiber but rather with stimulation sites along the next cochlear turn. It is possible that locations closest to the optical fiber might not be stimulated, because the spot size is small and the optical path may not include spiral ganglion cells next to the tip of the optical fiber.
In a realization of INS for a neural interface of the cochlea changing, the wavelength of the radiation can control the possible crossturn stimulation. The optical penetration depth of water, the main absorber involved in INS, is wavelength dependent. For the present experiments, the wavelength (1,869 nm) and consequently the penetration depth were kept constant. For radiation wavelengths between 1,830 and 1,900 nm, the water absorption curve is steep; that is, small changes in wavelength result in large changes in optical penetration depth. The effect of changing penetration depth on neural stimulation/recruitment can be seen in the change in optically evoked CAP amplitude (Izzo et al., 2008). The CAP amplitude increased with increasing penetration depth of the radiation, which was explained by an increasing number of neurons in the optical path being irradiated. Thus, it should be possible to employ small changes in wavelength to ensure that the desired cochlear structures are being irradiated.
The normal cochlea can encode sound pressure levels over a range of about 120 dB SPL. For electrical stimulation and conventional intrascalar stimulation paradigms, the dynamic range between the current amplitude at threshold and at maximum response is about 8–12 dB (Hartmann and Kral, 2004). More recently, a penetrating nerve array has been proposed as an alternative for electrical stimulation of the auditory nerve. A larger dynamic range of stimulation (13–17 dB) has been reported with this technique (Middlebrooks and Snyder, 2007). The results of the present experiments indicate that the dynamic range for optical stimulation is greater than that for contemporary cochlear electrical stimulation. However, the dynamic range for optical stimulation was limited in the present experiments by the stimulation threshold and the output power of the laser. In general, the stimulation threshold was at 15.3 μJ and the maximum energy per pulse (at 100 μs) from the laser was 127 μJ, which corresponds to a factor of 8.3 by which the energy could be increased. However, it is not clear whether electrical and optical values can be compared directly because optical energy is reported as a measure for INS, wheres current amplitude is used for electrical stimulation.
The more important question is whether optical stimulation indeed provides a greater number of discriminable level steps compared to cochlear electrical stimulation. This may be assessed by comparing each level with the next discriminable level step inducing a d′ change of 1. This approach has been applied by Middlebrooks and Snyder (2007). The number of stimulus levels between the cumulative d′ = 1 and d′ = 3 contours were measured at the electrode contact in the ICC that showed the lowest threshold. With those numbers, a “discrimination slope” was calculated. For electrical stimulation, this slope was expressed in units of d′ per decibel. For electrical stimulation with an intraneural electrode, the slope was 0.73 d′, for a monopolar and bipolar electrode configuration the values were 1.98 and 1.94 d′, respectively (Middlebrooks and Snyder, 2007). A similar calculation was performed for INS of the cochlea (Richter et al., 2011). When the slope is expressed in units of d′ per energy (μJ), it was on average 0.08 d′ (the average energy difference is 25.3 ± 21 μJ, N = 10). Expressed in , the slope was 0.42 d′ (the average dB value was 4.8 ± 2.3).
Although the results from acute studies demonstrate that optical stimulation may provide good dynamic range, we note that safety thresholds for long-term optical stimulation have not yet been defined. Photothermal interactions with the target or surrounding tissue at high radiant energies may alter tissue properties and reduce the dynamic range. Tolerance of optical stimulation at high radiant energies may also be a complicating factor. Long-term safety of optical stimulation and behavioral responses in awake animals need to be evaluated in the future.
The Mechanism of Stimulation
One of the key unanswered questions regarding IR neural stimulation is the mechanism of stimulation. Laser–tissue interactions can occur via various biophysical mechanisms, including photochemical, photothermal, photomechanical, and electric field effects (Jacques, 1992; Welch and van Gemert, 1995; Niemz, 2004). Wells and coworkers (Wells et al., 2006; Wells et al., 2007a) have investigated the various biophysical mechanisms that could be responsible for optical stimulation. Results from their study eliminated photochemical and electric field effects. Furthermore, they could not identify a single wavelength where stimulation was significantly enhanced, which would have suggested a photochemical mechanism. For photomechanical effects, stress waves, volumetric tissue, expansion and thermoelastic expansion were considered as mechanisms for INS. Wells et al. (Wells et al., 2006; Wells et al., 2007a) proposed that photomechanical effects from stress wave generation and thermoelastic expansion are unlikely means of depolarizing the neurons based on the finding that stimulation threshold for visible muscle twitch was independent of pulse duration (all tested at a nearly constant wavelength or depth of penetration). However, from the Wells et al. study (Wells et al., 2007a), the photomechanical effects from volumetric expansion could not be altogether excluded since a small, but not negligible, displacement of tissue did occur at radiant exposures slightly above threshold. Note that a pressure transient generated by a mechanical indenter and having the same temporal characteristics as the INS pulse failed to induce neural depolarization at pressures equivalent to 30× INS threshold. Furthermore, all IR pulse durations tested evoked an action potential. In a recent study, pressure waves resulting from the INS-induced volumetric expansion were investigated in more detail (Teudt et al., 2011) and have to be considered during stimulation of the auditory system (see discussions in [Teudt et al., 2011; Richter et al., 2011b]).
A photothermal mechanism was concluded to be the most likely means of depolarizing neurons via INS. A temporal and spatial thermal confinement is necessary to achieve optical stimulation (Wells et al., 2007a). Thermal confinement exists when the thermalized optical energy delivered by a single-pulse accumulates in the irradiated tissue before any of the heat can dissipate through conduction or convection. For thermal confinement to be achieved at these wavelengths, the pulse length of the stimuli is typically longer than 500 ns and shorter than 200 ms. The primary mechanism of INS is taken to be a transient temperature increase that results from absorption of the optical radiation by water in the tissue. Potential mechanistic candidates include thermal activation of a particular ion channel, thermally induced biophysical changes in the membrane (increase ion channel conductance, decrease in Nernst equilibrium potential), or through general expansion of the neural membrane that facilitates the flux of ions.
Results from a recent study to determine the mechanism of INS show that the rapid local increase in temperature with each laser pulse transiently alters the electrical capacitance of the plasma membrane, generating depolarizing currents in Xenopus laevis oocytes, mammalian cells, and artificial lipid bilayers (Shapiro et al., 2012). This latter mechanism is fully reversible and requires only the most basic properties of membranes common to all eukaryotes. It thus points to the potential generality of INS as a unique technology for control of excitable cells. The temperature-induced capacitive membrane current depolarizes the cell and the cell's inherent potpourri of ion channels, which then determines the characteristics for INS in the particular system.
Safety of Optical Stimulation
Any neuroprosthesis that will be implanted in patients is expected to remain in use over many years. For such applications, laser parameters that allow safe, chronic INS of the cells must be established. To date, limited data are available measuring safe stimulation over acute durations (Izzo et al., 2006b; Wells et al., 2007c; Rajguru et al., 2010). The present study explored various duration and radiant exposure levels during high repetition rate cochlear INS to evaluate injury thresholds. The cochlear function (CAP measurements) and histology after 3–5 hr of INS at various exposure levels were evaluated. Cochlear damage may arise from the propagation of large stress relaxation waves or from temperature increase. Laser-induced pressure waves in water are well documented for experiments with high local absorption (e.g., Welch and van Gemert, 1995; Niemz, 2004). In a recently published manuscript, it has been reported that pulsed 1,850-nm laser radiation, which is used for neural stimulation, also generates a measurable pressure (Teudt et al., 2011). At the maximum energy of the laser, which is coupled to a 200 μm diameter optical fiber, the peak-to-peak sound pressure can reach 62 dB SPL in air. As the cochlea is filled by endolymph and perilymph, it is of interest to what extent the laser-induced pressure waves exist when the absorbing volume is significantly decreased by immersion in water. Measurements in a swimming pool showed that radiant exposures of 0.35 J/cm2 generated a pressure of 31 mPa in water. Referenced to 1 μPa, this corresponds to a value of 89.9 dB, whereas referenced to 20 μPa the above pressure would be 63.8 dB SPL. Considering the approximate gain of 26 dB through the middle ear, a 63.8 dB SPL pressure in scala tympani would correspond to a sound level at the ear canal of 37.8 dB SPL at the maximum power of the laser. The same measurements were performed with radiant exposures at stimulation threshold for INS. The pressure in water was 0.35 mPa, which is 50.9 dB (re 1 μPa) and 24.9 dB SPL. With a transfer function of 1:20 for the middle ear, 24.9 dB SPL would correspond to ∼1.2 dB at the ear canal. As it relates to the data presented in this manuscript on the acute damage threshold for the guinea pig cochlea, it is not likely that the INS-induced pressure wave contributes to the damage. Rather, the acute damage is a result of the tissue heating.
For the acute cochlear damage, which is presented in this manuscript, the CAP amplitude was recorded because it is a sensitive marker for the physiological state of the cochlea. A small but nondamaging decrease in the temperature of the cochlea will decrease the CAP amplitude and will compromise cochlear function (Ohlemiller and Siegel, 1992, 1994). In the present experiments, the CAP amplitude remained stable over time if the pulse repetition rate was 250 Hz and laser energy was 25 μJ/pulse or lower. Furthermore, continuous stimulation time and the laser energy had an inverse relationship, that is the lower the stimulation energy was, the longer it was possible to stimulate without a functional loss. Each laser pulse that is delivered to the tissue results in a transient temperature increase. The heat is then dissipated by diffusion and convection. If the transient temperature increases occur quicker than the tissue can fully dissipate them, then the overall tissue temperature will increase with the potential risk of thermal damage. The net heating thus depends on the rate of laser pulses delivered, the temperature increase per pulse (proportional to the energy per pulse), and the speed by which the heat is distributed in the tissue. Dissipation and convection of the heat appear to be sufficiently fast in the guinea pig cochlea if the stimulation rate is 250 Hz and the radiant energy below 30 μJ/pulse. For larger cochleae, such as the cat or the human cochlea, this threshold energy for damage may be higher because the target volume that is available to distribute the heat is larger.
Cell death is characterized by the specific anatomical changes of the cell. Those changes may also provide insight in the mechanism for the cell death and can be classified according to its morphological appearance (which may be apoptotic, necrotic, autophagic, or associated with mitosis), its enzymological criteria (with and without the involvement of nucleases or of distinct classes of proteases, such as caspases, calpains, cathepsins, and transglutaminases), its functional aspects (programmed or accidental, physiological, or pathological) or its immunological characteristics (Thomsen, 1991; Pearce and Thomsen, 1995; Kroemer et al., 2009). Although markers of cell death may be seen, the process of cell death can be reversible until a “point-of-no-return” is reached. It has been proposed that the point-of-no-return is represented by massive caspase activation, loss of the membrane potential, complete permeabilization of the mitochondrial outer membrane, or exposure of phosphatidylserine residues (Kroemer et al., 2009).
In the present experiments, the time between the end of the stimulation and the harvest of the tissue was short. Cochlear sections were examined for obvious structural alterations of the tissue, edema, loss of the integrity of the cell's plasma membrane, cell shrinkage, complete fragmentation of the nucleus into discrete bodies, vacuole formation, and signs of coagulation. None of the latter signs could be seen in the cochlear cross-sections that were examined. Furthermore, the outer and inner hair cells, the tectorial membrane, and the basilar membrane could be identified in all sections obtained from cochleae that were exposed to radiant energy levels of 127 μJ/pulse, 100 μs pulse duration, and 250 Hz pulse repetition rate. Thus, the functional measurements of cochlear CAP were a more sensitive method to assess any damage owing to the laser irradiation.
INS bears the risk of thermal damage to the tissue from the instantaneous temperature increase or from potential accumulation of thermal changes. The immediate temperature increase depends on the energy of the pulse, and the temperature accumulation on the rate of the pulses is presented. With the present study, we have shown that the injury thresholds in guinea pig cochleae for acute INS is between 25 and 30 μJ/pulse at a radiation wavelength of 1,869 nm, for 100 μs laser pulses, which are delivered at 250 Hz. The results reflect short-term results and long-term damage that may occur after weeks of stimulation cannot be excluded with the present experiments.
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