Electrical stimulation of nerves has been demonstrated as a viable means of neuroprosthesis. Peripheral neuroprostheses typically use interfaces with peripheral nerves or muscles to restore motor and sensory functions [1
]. Of these interfaces, the Utah Slant Electode Array (USEA) has been shown to provide comprehensive access to multiple independent motoneuron subpopulations and to effect local stimulation with low current levels, even in a long-term implant [2
]. These multielectrode arrays are designed to penetrate deeply into neural tissue, for example, to access the different fascicles within the peripheral nerve, as shown in . A typical USEA can reach up to 1.5-mm deep into tissue, but electrodes as long as 9 mm have been machined using alternative fabrication methods [6
]. The variable length intrafascicular electrodes facilitate low stimulation threshold and selectivity with a few electrodes targeting a specific muscle [2
]. The electrodes are placed in proximity to the target axons and thus bypass the epineurium and perineurium, which act as insulators that make stimulation from around the nerve (e.g., via cuff electrodes) weaker and less selective. In addition, the high-channel count (i.e., 100 electrodes) allows successive recruitment of neurons to evoke maximal fatigue-resistant forces in various muscles [7
]; graded activation of several muscles with the USEA for normal multi-joint motions was demonstrated by producing a graceful feline stance with paralyzed hind limbs of the cat [4
Fig. 1 Utah Slant Electrode/Optrode Array for peripheral nerve stimulation and/or recording. (a) Transverse cross-section of cat sciatic nerve with single row of slant array shown. The microneedles penetrate through epineurium, perineurium and endoneurium to (more ...)
Electrical signals, on the other hand, cause stimulus artifacts that prevent simultaneous recording and stimulation of adjacent neurons. An alternative modality using infrared light as input energy avoids the artifact problem due to the absence of direct charge injection [9
]. In the initial study [9
], a pulsed IR laser was coupled to a 600-μ
m diameter optical fiber and the light was incident on a single spot on a rat sciatic nerve surface. Action potentials were selectively evoked without causing histological tissue damage. Stimulation and ablation fluence thresholds were determined for several wavelengths between 2.1 and 6.1μ
m. Optimal wavelengths with a safety margin between stimulation and damage threshold for extraneural stimulation of mammalian peripheral nerves was observed to depend on the absorption spectrum of water, the dominant tissue absorber in the IR; damage threshold was strongly affected by the absorption, but stimulation commencement levels varied less. Stimulation at 3 and 6-μ
m wavelengths, which have very high absorption (i.e., approximately zero penetration depth) as shown in , readily created nerve damage with threshold stimulation fluence; 2.1 and 4-μ
m wavelengths, which have smaller absorption coefficients, were found to be particularly well suited for infrared neural stimulation (INS). This finding, along with recordings of surface nerve temperatures, suggest that INS acts in nerve via the induction of a spatio-temporal heat gradient in the tissue (4°C at the axonal level) with about 3-mJ/mm2
extraneural threshold dose [10
]. Although the underlying physiological mechanism is still under investigation, the threshold optical dose closely follows the water absorption spectrum. A recent study however, determined that IR is absorbed by water during INS and causes local tissue heating that depolarizes the target cell by changing the membrane electrical capacitance [11
Water absorption curve for IR wavelengths, which is representative of tissue absorption in the IR. 1.87 μm is recommended for peripheral nerve INS, but 2.1 μm has also been extensively used.
With a 2.12-μ
m (Ho:YAG laser) input, INS has been demonstrated to achieve a selective excitation volume with respect to extraneural electrical stimulation [12
]. A wavelength of 2.12 μ
m has been extensively used because it causes minimal nerve damage and can be generated from a commercially available Ho:YAG laser that is currently utilized in many clinical applications [9
]. This wavelength corresponds to a tissue penetration depth between 300 and 500 μ
m as supported by Fourier transform infrared spectroscopy results in [12
], which is deemed suitable for stimulation based on rat peripheral nerve geometry [13
]; a wavelength in the vicinity of 1.87μ
m, having similar absorption characteristics as 2.12 μ
m, has also been shown to stimulate effectively in various applications [14
]. However, neural tissue of different types and morphologies may require different wavelengths for optimal stimulation such that the optical penetration depth is matched to the targeted excitable tissue.
As with electrical stimulation, an intrafascicular multiple access approach for INS will likely provide coverage of a large number of independent neuron subpopulations, lower activation energy, and better spatial selectivity than extraneural INS. These advantages may be demonstrated by using optrode arrays made from intrinsic silicon (Si), with IR light coupled from a pulsed or continuous wave (CW) laser source operating at wavelengths from 1.1 to 5.5μm, as indicated in . Ultimately, INS with an optrode array is expected to yield a greater separation between stimulation and damage thresholds and permit a wider range of wavelengths to effect a neural response - light does not need to travel through the connective tissues within the nerve in order to reach the axons, potentially allowing wavelengths with high absorption to be used.
Other penetrating probes for optical stimulation have been used in the field of optogenetics, where visible light is delivered to excite genetically targeted neurons expressing light-sensitive channels (e.g., ChR2). Tapered optical fibers serving as tissue-penetrating optical probes have been utilized [18
]. A single optrode made of a 50-μ
m multi-mode fiber was inserted in mouse brain slices to trigger localized epileptiform events in a single cortical site [18
]. Simultaneous delivery of visible light to multiple neuronal sites have been achieved by arranging tapered single-mode optical fibers in a 2D array of optrodes [20
]; etched fibers were glued to commercially available planar silicon probes, which are an alternative to tungsten electrodes utilized in previous studies [23
]. Experiments with ChR2-transfected rat hippocampus demonstrated multiple local stimulation via these 2D arrays. A more sophisticated microfabricated 2D multiwaveguide probe was introduced as an alternative to crude arrays of tapered fibers [24
]. The probe targets points along its axis (i.e., stimulation of sites along the depth rather than the lateral direction) using parallel independent single-mode rectangular waveguides of silicon oxynitride core and oxide cladding that converge into a probe structure. Each waveguide ends at a particular target depth with a corner aluminum mirror to perform side-firing; transmission efficiency ranged from 23 to 33% as determined from bench testing.
We have recently developed a micromachined 3D optrode array for infrared neural stimulation. This array covers a wide area of neuronal stimulation sites and reaches targets at varying depths, thereby facilitating high-channel-count optical stimulation. Preliminary optical [25
] and intrafascicular
] testing results have been reported. In this paper, we perform detailed characterization of this early-generation Utah Slant Optrode Array (USOA) neural interface. The USOA is designed after the USEA architecture, which has been adapted for numerous physiological requirements, such as nerve dimensions and axon depth, through variation in electrode length and spacing. Note that the same design can be adapted to other transparent substrates for both INS and optogenetic applications.