An implantable wireless tactile sensing system

The sense of touch is critical to dexterous use of the hands and thus an essential component to efforts to restore hand function after amputation or paralysis. Prosthetic systems have focused on wearable tactile sensors. But wearable sensors are suboptimal for neuroprosthetic systems designed to reanimate a patient’s own paralyzed hand. Here, we developed an implantable tactile sensing system intended for subdermal placement. The system is composed of a microfabricated capacitive force sensor, a custom integrated circuit supporting wireless powering and data transmission, and a laser-fused hermetic silica package. The miniature device was validated through simulations, benchtop testing, and ex vivo testing in a primate hand. The sensor implanted in the fingertip accurately measured skin forces with a resolution of 4.3 mN. The output from this novel sensor could be encoded in the brain with microstimulation to provide tactile feedback. More broadly, the materials, system design, and fabrication approach establish new foundational capabilities for various applications of implantable sensing systems.


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Skin is endowed with a variety of tactile mechanoreceptors that play an invaluable role in sensing our 48 environment and guiding movement. A lack of tactile feedback severely limits dexterity of both biological 49 (1) and biomimetic robotic systems (2). Thus, multiple approaches have been used to realize artificial 50 tactile sensors for wearable electronics, prosthetics, and robotics (3-5). In general, these tactile sensors 51 can be classified according to the sensing principles employed, such as capacitive (6-9), piezoresistive 52 (10-17), and magnetic field-based sensing technologies (18,19). These sensors have been embedded in 53 flexible and stretchable substrates for applications such as healthcare monitors (20), prosthetic skins (21), 54 patient rehabilitation (22), electronic skins (6), and robotic skins (23). Wearable designs such as sensorized 55 gloves or skin-attached electronics are commonly adopted for these sensor systems. 56 However, wearable sensors are suboptimal for applications in which the goal is to restore a sense of 57 touch to biological skin, as in the case of paralysis. Paralysis disrupts both motor and somatosensory 58 signals between the brain and body. Brain-machine interface (BMI) technology has been used to 59 implement brain-controlled muscle stimulation to reanimate a paralyzed limb (24,25). This strategy can 60 restore volitional hand movement in humans with tetraplegia (26, 27). But tactile sensing in the palm and 61 fingertips that is critical for dexterous manipulation is still lacking. One could consider augmenting such 62 systems with tactile sensors coupled with appropriate neural stimulation to artificially encode the sense of 63 touch (28). Although wearable sensors are a potential candidate for supplying this tactile functionality 64 noninvasively, they have several disadvantages. Skin-attached devices have limited longevity due to 65 epidermal turnover and environmental interference. Wearables place often ill-fitting material between the 66 skin and grasped object, thus altering the natural interface. Finally, bulky sensorized gloves could place 67 undue burden on activated muscles already prone to fatigue due to unnatural recruitment (29). 68 An alternative approach to resolve these issues is an implantable tactile sensor system. Implantable 69 microelectromechanical systems (MEMS) sensors and actuators have been adopted to monitor human 70 health conditions, improve quality of life, and save lives (30-32). However, long-term implantable 71 technology is extremely challenging due to the requirements on hermeticity, biocompatibility, as well as 72 a proper form factor (33). The system should attain long-term hermeticity to provide any enclosed 73 electronic circuitry with protection from the harsh environment of the human body. Simultaneously, the 74 packaging material should be biocompatible and small enough to fit within the target location; and ideally 75 should also support communication means so data (and optionally power) can be transferred to and from 76 the sensor.  The operation principle of the capacitive sensor is illustrated in Figure 2. When there is no force applied 150 to the sensor, the upper silica plate, with a floating potential upper electrode, is parallel with the middle 151 silica plate, containing two lower electrodes ( Fig. 2A) will continue to deform and conform to the lower plate electrodes.

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To quantitatively evaluate the sensor operation principle, a COMSOL Multiphysics simulation using 160 the solid mechanics module and the AC/DC electrostatics module was performed (Fig. 2D). The loading 161 procedure was simulated in a two-step fashion (Fig. S2). In the first step, the solid mechanics module was    Hz sinusoidal force ranging from 2 to 7 N resulted in sinusoidal change in capacitance in the range of 7.2 221 pF to 8.2 pF (Fig. 3D). These values agreed well with the static force output capacitance of 7.0 pF at 2 N 222 and 8.1 pF at 7 N.

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Another important mode of operation for in vivo use is to sense hydrostatic pressure rather than direct 224 compressive forces on the sensing membrane. To understand this mode of operation, the sensor was placed 225 in a water-filled syringe connected to a pressure gauge (Fig. 3E). The capacitance of the sensor inside the 226 syringe increased from 6.4 pF to 10.5 pF when pressure increases from 2.3 psi to 45 psi, which corresponds 227 to 0.2 N to 3.9 N considering the dimension of the sensing membrane (Fig. 3F) built, force-controlled motorized stage was used to apply small forces to the fingertip, in the range of 239 physiological light touch (< 1N), while wirelessly recording the sensor response from above the fingernail 240 (Fig. 4A). The sensor response to static forces was repeatable and nearly linear, with a sensitivity of 0.8 241 pF/N (Fig. 4B).  Time-varying forces were subsequently applied through manual indentation of the fingertip with a load 252 cell recording applied forces (Fig. 4C). The sensor capacitance changes measured wirelessly by an external 253 base unit closely followed the dynamic forces (Fig. 4D). A simple linear transformation of the sensor 254 output could reliably estimate (R 2 = 0.94) the applied dynamic forces (Fig. 4E). Together, the results show 255 that the capacitive sensor system has a form factor suitable for subdermal implantation in the fingertip and 256 can be used to faithfully estimate tactile forces for use in sensory restoration systems.

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Discussion 259 We developed a novel wireless capacitive force sensor intended for long-term operation within the body.

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Force transduction occurred through mechanical deformation of a hermetic fused silica package housing unit. The device was validated through simulation, benchtop testing, and ex vivo testing in a primate hand.

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The sensor microsystem was designed to function as an artificial mechanoreceptor for neuroprosthetic

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The Ti layer is patterned using lift-off to complete the substrate fabrication.

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Fabrication of the lower plate 370 To provide a hermetic encapsulation of the electronics, a bottom plate with deep cavity of 600 μm is 371 required. A two-step CO2 laser fabrication is adopted for the process (Fig. S1). To fabricate the cavity in was connected to an oscilloscope (Fig. S4). To approximate implanted conditions, the Bose test instrument 418 was fitted with a flat surface acrylic loading component (10 mm × 10 mm square shape) that applied forces 419 to a silicone layer (Ecoflex0030, 5 mm in thickness and 5 mm in diameter) overlying the sensor upper 420 plate (Fig. 3A). Wireless sensor operation was the same as described above. In the presence of normal 421 forces, both the silicone and sensor upper plate deformed.

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Hydrostatic pressure experiments 424 The sensing system was placed inside a 5 ml syringe which was filled with water (Fig. 3E). The base unit 425 was set outside the syringe for wirelessly powering the sensing system and collecting data from the system.

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A digital pressure gage (MG1, SSI Technologies) was connected to the syringe to measure pressure as the 427 syringe plunger was manually depressed to achieve target static pressures.

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Ex vivo experiments 430 The right hand of a rhesus macaque (Macaca mulatta) that had recently been euthanized for clinical