Biomed Microdevices (2015) 17:37 DOI 10.1007/s10544-014-9923-8

Microfabricated infuse-withdraw micropump component for an integrated inner-ear drug-delivery platform Vishal Tandon · Woo Seok Kang · Abigail J. Spencer · Ernest S. Kim · Erin E. L. Pararas · Michael J. McKenna · Sharon G. Kujawa · Mark J. Mescher · Jason Fiering · William F. Sewell · Jeffrey T. Borenstein

© Springer Science+Business Media New York 2015

Abstract One of the major challenges in treatment of auditory disorders is that many therapeutic compounds are toxic when delivered systemically. Local intracochlear delivery methods are becoming critical in emerging treatments and in drug discovery. Direct infusion via cochleostomy, in particular, is attractive from a pharmacokinetics standpoint, as there is potential for the kinetics of delivery to be well-controlled. Direct infusion is compatible with a large number of drug types, including large, complex molecules such as proteins and unstable molecules such as siRNA. In addition, hair-cell regeneration therapy will likely require long-term delivery of a timed series of agents. This presents unknown risks associated with increasing the volume of fluid within the cochlea and mechanical damage caused during delivery. There are three key requirements for an intracochlear drug delivery system: (1) a high degree of miniaturization (2) a method for pumping precise and small volumes of fluid into the cochlea in a highly controlled manner, and (3) a method for removing excess fluid from the limited cochlear fluid space. To that end, our group is developing

a head-mounted microfluidics-based system for long-term intracochlear drug delivery. We utilize guinea pig animal models for development and demonstration of the device. Central to the system is an infuse-withdraw micropump component that, unlike previous micropump-based systems, has fully integrated drug and fluid storage compartments. Here we characterize the infuse-withdraw capabilities of our micropump, and show experimental results that demonstrate direct drug infusion via cochleostomy in animal models. We utilized DNQX, a glutamate receptor antagonist that suppresses CAPs, as a test drug. We monitored the frequency-dependent changes in auditory nerve CAPs during drug infusion, and observed CAP suppression consistent with the expected drug transport path based on the geometry and tonotopic organization of the cochlea. Keywords Drug Delivery · Microfluidics · Inner Ear · Intracochlear · CAP · DPOAE · Micropump

1 Introduction Vishal Tandon and Woo Seok Kang contributed equally to this manuscript.

1.1 Background

V. Tandon · A. J. Spencer · E. S. Kim · E. E. L. Pararas · M. J. Mescher · J. Fiering · J. T. Borenstein () Charles Stark Draper Laboratory, Cambridge, MA 02139, USA e-mail: [email protected]

360 million people worldwide have disabling hearing loss, as estimated by the World Health Organization in 2014 (http://www.who.int/pbd/deafness/en), and treatments represent a $10 billion potential market (Vio and Holme 2005) . Treatment of severe auditory disorders is challenging because (i) identification of drug targets is non-trivial owing to the complexity of the auditory pathway, and (ii) the limited accessibility of the inner ear hinders targeted, controlled delivery of potentially therapeutic compounds. Recent advances in the molecular biology of hearing (Holley 2005; G´el´eoc and Holt 2014) have identified drug

V. Tandon · W. S. Kang · M. J. McKenna · S. G. Kujawa · W. F. Sewell Department of Otology and Laryngology Massachusetts Eye and Ear Infirmary and Harvard Medical School Boston, Boston, MA 02114, USA W. F. Sewell e-mail: William [email protected]

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targets and therapeutic compounds that have the potential for both ototoxicity protection (Li et al. 2001; Rybak and Whitworth 2005) and treatment of debilitating hearing disorders, including sensorineural hearing loss (SNHL) (Izumikawa et al. 2005; Mizutari et al. 2013), noiseinduced hearing loss (NIHL) (Lynch and Kil 2005), and tinnitus (Murai et al. 1992; Shulman and Goldstein 2000; Eggermont 2005). One of the major challenges in treatment of auditory disorders is that many of the relevant drugs and drug candidates are toxic or have severe side effects when delivered systemically. For example, steroids are commonly used for treatment of SNHL (Wilson et al. 1980; Chandrasekhar et al. 2000; Haynes et al. 2007), but their systemic administration has been associated with hyperglycemia, hypertension, hypokalemia, peptic ulcer disease, osteoporosis, immunosuppression, and adrenal suppression (Chrousos 2007; Garc´ıa-Berrocal et al. 2008; McCall et al. 2010). Inhibition of Notch signaling with a γ -secretase inhibitor has shown potential for hair-cell regeneration, but γ -secretase inhibitors are toxic when delivered systemically (Mizutari et al. 2013). In addition, systemic delivery is often ineffective due to insufficient drug concentration at the target site caused by a combination of general degradation/dilution of the drug (reducing its effective concentration), and the inability of the drug to cross the blood-cochlear barrier (Haynes et al. 2007; Juhn and Rybak 1981; Juhn 1988; Juhn et al. 2001). Local delivery to the inner ear is expected to allow more precise dose control, intensify therapeutic effects, and minimize adverse side effects. One of the most common approaches to local delivery is intratympanic injection of drug onto the round window membrane (RWM) (Doyle et al. 2004; Haynes et al. 2007), often with the aid of a biodegradable gel (Endo et al. 2005; Ito et al. 2005; Iwai et al. 2006), or alginate beads (Noushi et al. 2005). RWM delivery of small-molecule drugs, however, relies on passive diffusion of drug across the membrane. The kinetics of this transport process are complex (Salt and Plontke 2005) and often unpredictable, making RWM delivery less attractive for pharmacokinetic studies. Furthermore, there is growing interest in therapies requiring inner-ear delivery of large, complex molecules or molecules with limited stability, such as growth factor proteins (Richardson et al.2004; Bianchi and Raz 2004; Richardson et al. 2006, 2009), siRNA (Maeda et al. 2009), and vectors for gene therapy (Praetorius et al. 2007); RWM delivery can be more challenging or intractable with these classes of molecules. In cases where precise delivery kinetics are required and/or large molecules need to be delivered, direct intracochlear infusion via cochleostomy or penetration of the RWM is preferred (Himeno et al. 2002; Chen et al. 2005, 2006). Direct infusion is more invasive, but offers

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advantages of highly targeted delivery, lower drug doses, minimized systemic toxicity, and controllable delivery kinetics. Delivery can be controlled with the aid of a micropump that maintains a set flow rate for drug infusion. Osmotic pumps have been commonly used in pre-clinical studies of drug delivery via cochleostomy (Swan et al. 2008), but flow control is limited with such pumps, and their total payload may be insufficient for some treatment courses. Progress toward hair-cell regeneration through differentiation of progenitor cells (Li et al. 2003; 2004) suggests that clinically relevant treatments will likely require highly localized delivery of a timed series of reagents over a period ranging from months to years (Sewell et al. 2009). This suggests that hair-cell regeneration treatment will require a micropump that is wearable and/or implantable and that can maintain delivery of precise doses of drug(s) over extended time periods. One of the challenges with direct-infusion drug delivery is the relatively unknown potential for causing mechanical damage to the delicate structures in the organ of Corti, either from high shear stresses or large infusion volumes leading to increases in the hydrostatic pressure of the scala tympani. Fridberger et al. (1997) showed that pressure changes on the order of 10 Pa can lead to shifting of the basilar membrane position. In order to mitigate the increases in hydrostatic pressure accompanied by an increase in the volume of fluid within the scala tympani, our group has designed systems for reciprocating infusion—infusion of concentrated drug into the scala tympani followed by withdrawal of perilymph with a lower drug concentration at a later time, resulting in net delivery of drug with zero net increase in the volume of fluid in the scala tympani (Chen et al. 2005; Fiering et al. 2009; Sewell et al. 2009; Pararas et al. 2011; Kim et al. 2014). These systems utilized microfabricated fluidic components (e.g. drug reservoirs, fluidic capacitors) combined with commercially available micropumps and actuation systems. For long-term infusions, these components were packaged into a wearable headmount format for guinea pig animal models (Fiering et al. 2009; Kim et al. 2014).

Fig. 1 Photograph of our electromagnetic-actuator-based micropump with microfluidics, PCB, and actuator fixture

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Here we present a design for a miniaturized micropump (Fig. 1) that incorporates all of the essential fluidic elements into a single, integrated device that is microfabricated from machined, laminated stacks of polyimide layers (Mescher et al. 2009). Flow is actuated and controlled through the use of custom-machined, miniature electromagnetic actuators that are mounted to the fluidic components. The device is designed to be used as part of a long-term drug-delivery system that is head-mounted to guinea pig animal models, like in our previous generations of devices (Fiering et al. 2009; Kim et al. 2014). The integrated design has several important advantages: (i) increased miniaturization, (ii) increased robustness due to fewer fluidic connections between components, (iii) simplified design and easier operation due to the reduction in the overall number of components, (iv) more control over the specifications of the pump and fluidic components, because all components are custom-built, (v) independent control over infuse flow rates, withdraw flow rates, and delays between them—withdrawal of perilymph does not have to immediately follow infusion of drug, as was the case for the system designed by Kim et al. (2014), and (vi) the potential for automated delivery of multiple reagents. As a proof of concept, we fabricated a simplified version of the pump in order to demonstrate its infuse-withdraw capabilities. We demonstrated reciprocating delivery with bench-top measurements of the flow rates generated by our micropump. We also demonstrated infuse-only drug delivery in an animal model. For this, we used our micropump to infuse a glutamate receptor antagonist, DNQX, into guinea pig cochleae via cochleostomy at the basal turn. DNQX disrupts hair-cell-to-auditory-nerve synaptic transmission, and therefore suppresses auditory-nerve compound action potentials (CAPs) (Littman et al. 1989; Sewell 1996; Parks 2000). We measured CAPs generated in response to tone-pip stimuli at characteristic frequencies ranging from 2.78–32 kHz as functions of time after the start of DNQX delivery to the cochlea. From the well-known tonotopic organization of the cochlea (Greenwood 1996), we inferred the distribution of drug over time from the frequency-dependent suppression of the CAPs. Concurrent measurements of distortion product otoacoustic emissions (DPOAEs) during infusion were used to ensure that no other damage to middle/inner ear structures was caused by our infusion system.

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Fig. 2 Scheme of a 4–port micropump that has a central pump chamber that is fluidically connected to 4 valves, and an integrated drug reservoir. Each valve corresponds to a port, and is numbered

of fluid that can be pumped into or out of the device, ii) an outlet for infusing and withdrawing fluids from the drug-delivery target, iii) a drug reservoir inlet, and iv) a drug reservoir outlet. Each of the four ports is individually addressed with a valve, and the valves are fluidically connected to a central pump chamber. Coordinated closing and opening of any two valves along with actuation of the pump chamber leads to net flow through the associated ports. The delivery process is as follows: i) endogenous fluid is withdrawn by the pump into the fluidic capacitor (ports 1 and 2); ii) drug is loaded into the pump chamber (ports 3 and 4); iii) a mixture of drug and endogenous fluid is infused—this can be done with multiple, timed infusions of small boluses (ports 1 and 2); iv) at a later time, endogenous fluid is withdrawn into the fluidic capacitor again (ports 1 and 2). The drug loading process is similar to that presented by Kim et al. (2014). As a first step, we fabricated a simplified version of the 4–port micropump design (Fig. 3) that omitted the internal drug reservoir and the 2 associated valves (ports 3 and 4

1.2 System overview In contrast to most commercial micropumps, which only support unidirectional pumping from an inlet to an outlet (2 ports), our micropump design enables bidirectional pumping and can support pumping to and from multiple ports. Our goal is to fabricate a pump with 4 ports (Fig. 2) that are connected to: i) a fluidic capacitor for on-device storage

Fig. 3 Photograph of the fluidics portion of the micropump, which comprises several layers of polyimide films (Kapton) that are laminated together

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in Fig. 2). The simplified version of the pump was sufficient for bench-top characterization of its ability to generate reciprocating flow. For the animal experiments, an external drug reservoir was connected to the inlet of the micropump, and the pump was operated in infuse-only mode, because drug loading into endogenous fluid as described above cannot be achieved without the internal drug reservoir and associated valves. The ceiling of the the pump, valve, and capacitor fluid chambers is a thin (25.4 μm), flexible polyimide membrane. The membrane can be deflected to block flow in the valves, or to displace fluid in the pump chamber in order to actuate flow. As described previously by Mescher et al. (2009), we also utilize the compliance of the membrane in the integrated fluidic capacitor for fluid storage. All of the straight microfluidic channels in the device have a width of 384 μm. The capacitor chamber has a diameter of 14 mm, and a height of 0.25 mm (nominal, when the fluid pressure inside the capacitor is at equilibrium with the pressure outside of the capacitor). The pump chamber has a diameter of 4 mm and a nominal height of 0.25 mm. The annular valve-seat channels have outer diameters of 3.1 mm and inner diameters of 1.5 mm. The valve-seat through holes have diameters of 0.7 mm.

2 Materials and methods

in this study, was composed of, in mM: NaCl, 120; KCl, 3.5; CaCl2 , 1.5; glucose, 5.5; HEPES, 20. NaOH was used to adjust the pH of the AP to 7.5 (resulting in a total Na+ concentration of 130 mM). DNQX was dissolved in artificial perilymph solution at a concentration of 300 μM. 2.2 Fabrication Detailed descriptions of fabrication of microfluidic devices from polyimide films have been described previously by Mescher et al. (2009). Briefly, features of the micropump were designed using SolidWorks CAD software and converted for use with machining equipment using inhouse written MATLAB scripts. Microfluidics features were machined into polyimide layers 1–5 (Fig. 4) using a Quickcircuit 7000 XY router table (T–Tech, Norcross, GA). In layer 4, only the large (4–mm and 14–mm diameter) holes were machined on the router. The remaining features, including fluid channels that were cut partially to a depth of 100 μm, as well as valve through holes, were machined using a UV excimer laser (IX–1000, IPG Photonics, Oxford, MA) after layers 3 and 4 were laminated together (described in the next paragraph). After laser machining, the layers were cleaned in oxygen plasma (PX–250 Plasma Chamber, March Instruments, Concord, MA) for 5 min at a power of 150 mW and a pressure of 280 mTorr in an oxygen-nitrogen atmosphere with 30 % O2 .

2.1 Materials and reagents Polyimide (Kapton) sheets with thicknesses of 1 mil (25.4 μm), 3 mil (76.2 μm), and 5 mil (127 μm), were purchased from Fralock (Valencia, CA). 0.5 mil (12.7 μm)– thick R/flex 1000 bonding films were obtained from Rogers Corporation (Chandler, AZ). Double-sided, copper-clad (1 oz.), 1.57–mm thick prototyping board was obtained from Digi-Key (Thief River Falls, MN). 28–Ga 316 stainless-steel tubing (178 μm ID) was purchased from McMaster Carr (Robbinsville, NJ). 360–μm OD polyether-ether ketone (PEEK) tubing (with IDs of 50 μm, 75 μm, and 150 μm) and Tygon tubing (0.25– mm ID/0.76–mm OD and 1.3–mm ID/3.02–mm OD) were obtained from IDEX Health & Science (Oak Harbor, WA). Polytetrafluoroethylene (PTFE) tubing with an ID of 101 μm and an OD of 201 μm was purchased from Zeus Inc. (Sub Lite Tubing, Branchburg NJ). The PTFE tubing was treated with Fluoroetch (Acton Technologies, Pittston, PA) according to the manufacturer’s instructions in order to increase its ability to be bonded with adhesives. 6,7–dinitroquinoxaline–2,3–dione (DNQX) was purchased from Fisher Scientific (Waltham, MA). All other chemicals and reagents were purchased from Sigma-Aldrich (St. Louis, MO). Artificial perilymph (AP), used as a control

Fig. 4 Scheme of the polyimide, PCB, and 3D-printed actuator fixture layers that make up the 2-port micropump. The numbered layers were fabricated from polyimide, and their thicknesses are indicated

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The polyimide layers were laminated together under high temperature and pressure (175◦ C, ramped at 5◦ C/min, and then held for 1 hour, and 2000 kPa effective pressure) using interspersed R/flex 1000 bonding films. Lamination was carried out in 3 steps, where first layers 1 and 2 were laminated together, then layers 3 and 4 were laminated together (and laser machined), and finally the entire polyimide stack was laminated together (Fig. 3). During the lamination process, the 25–μm polyimide layer was clamped into a custom-machined annular clamp, so that it would remain taught. After lamination, a 1.5-cm length of stainless-steel tubing was inserted into the inlet microfluidic channel of the polyimide stack, and another identical tube was inserted into the outlet. The tubes were fixed in place with epoxy (Epoxy 907, Miller-Stephenson, Danbury, CT). A printed circuit board (PCB) for the actuators was routed from double-sided prototyping board using the Quickcircuit 7000 router. A Samtec ZF5S 10–pin, 0.5–mm pitch zero-insertion-force connector (Samtec, New Albany, Indiana) was soldered onto the PCB, and used to make the connection between the actuator PCB and a custom-built controller board via a Samtec FJH 0.5 mm flat flexible cable. A NI USB–6501 Digital I/O device (National Instruments, Austin, TX) was used as an interface between the controller board and a PC running an in-house-written LabView code that was used to control the actuator pumping sequences. The actuator fixture and cover layers were 3D–printed on a Viper Si2 SLA stereolithograpy system using Accura60 plastic as the working material (3D Systems, Rock Hill, SC).

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the washer and the magnet. The button is pressed against the membrane fluidic layer by a spring that passes through the center of the annular electromagnet (Fig. 5a). In this state, if the actuator is controlling a valve, the membrane is deflected so that it blocks the microfluidic flow path. When the electromagnet is supplied power, the washer is attracted to it, and the button head retracts from the membrane (Fig. 5b). This creates a path for fluid flow, and corresponds to opening the valve. Cycling an actuator between the powered and unpowered states results in bursts of flow, but no net flow through the device. An actuator situated above a fluidic chamber, with no valve, can create larger bursts of flow. In our device, such a displacement chamber is located at the center, and is fluidically connected to each of the other valves (2 for this device, 4 for the concept shown in Fig. 2). Coordinated opening and closing of any 2 valves along with actuation of the displacement chamber results in net flow. Different pumping schemes (Table 1) enable control of the direction and magnitude of flow.

2.3 Electromagnetic actuators Annular electromagnets that were custom-machined inhouse, in combination with springs and ferromagnetic washers, are used as actuators in our micropump. The electromagnets have outer diameters of 7.6 mm, inner diameters of 3.2 mm, and are 1.8 mm in height. The washers have outer diameters of 7.2 mm, inner diameters of 2.55 mm, and are 0.4 mm in height. Polycarbonate snap buttons are inserted into the centers of the washers, and they extend 0.764 mm beyond the surfaces of the washers. The electromagnets are actuated by applying a 12–V spike voltage for 2 ms followed by a 2–V hold voltage for 300 ms. The actuator assemblies are surface-mounted to the membrane side of the microfluidic layers and are supplied current through the actuator PCB. The actuators are held in place by countersinks in the PCB, and by the 3D–printed fixture and cover structures (Fig. 4). The microfluidcs, PCB, and 3D–printed structures are clamped together with bolts. When an actuator is unpowered, a washer that has a polycarbonate button head sits below it with a gap between

Fig. 5 Scheme of an electromagnetic actuator used for valve control and pumping in its (a) unpowered and (b) powered state

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Table 1 Summary of pump schemes

“I” refers to the inlet valve actuator, “O” refers to the outlet valve actuator, and “P” refers to the pump actuator. The times listed correspond to the amount of time the indicated actuators are powered (held at 2 V). When an actuator is switched from unpowered to powered, a 2–ms, 12–V spike precedes the hold voltage. “W” refers to the wait condition, where no current is supplied to any actuator. Legend: ♦ Closed Valve, Unpowered Valve Actuator; Open Valve, Powered Valve Actuator;  Unpowered Pump Actuator; *with 6 cm of 75-μ m ID PEEK tubing attached to the outlet.

2.4 Device characterization The performance of our micropump was characterized in terms of valve leak rates and flow rate output. Valve leak rates were characterized by applying 20.7 kPa of water pressure at the inlet via a pressure controller (Model No: PCD–15PSIG–D–PC, Alicat Scientific, Tucson, AZ), and inferring the amount of water leaving the outlet by measuring the distance of travel of the air-water interface within a 0.25–mm ID Tygon tube attached to the outlet over a time period of 12 min. In order to measure the leak rate of the inlet valve, the pump and outlet actuators were held in the “powered” state by permanent magnets placed on top of the device, just above the actuators. Similarly, for measurement of the outlet valve, the pump and inlet actuators were held in the “powered” state. Permanent magnets were used for these tests because a continuous supply of current to an electromagnet actuator can lead to overheating and damage. During normal device operation, valves do not typically need to be held open. The flow rate output of the micropump depends on the applied load and on the pumping scheme used (Table 1). Pumping schemes described here were chosen for applications in intracochlear drug delivery. In order to allow for comparisons between different pumping schemes, flow rates are reported as volume per pump stroke, or stroke volume. A measured flow rate can be converted into stroke volume by dividing by the stroke frequency (number of

Powered Pump Actuator;

Capacitor.

pump strokes per second). The average pump flow rate was measured by first priming the device with water, and then connecting the inlet to a water reservoir via a water-filled, 0.25–mm ID Tygon tube. A 1.3–mm ID Tygon tube (initially filled with air) was used to connect the outlet of the pump to a pressure source. Pump scheme 1 (fast infuse) was then activated for 60 s, and the distance the air-water interface was displaced in the outlet tube was measured in order to infer the flow rate. Instantaneous flow rates were measured by connecting a Sensirion SL 1430 Liquid Flow Sensor (Sensirion, St¨afa, Switzerland) to the outlet of the micropump. Flow rates were recorded using the included Sensiview software at a sampling rate of 50 Hz. This method could not be used unless a load or fluidic resistor was applied at the outlet of the pump, because at zero or small loads, the instantaneous flow rates generated by the micropump exceeded the flow sensor’s maximum range of 40 μL/min. 6 cm of 75–μm ID PEEK tubing was used as a fluidic resistor in both pump characterization and drug infusion experiments. 2.5 Surgical procedures All procedures used in this study were approved by the Massachusetts Eye and Ear Infirmary Institutional Animal Care and Use Committee. Detailed descriptions of surgical procedures and hearing tests have been described previously by (Chen et al. 2005). Briefly, four male albino guinea pigs

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(Hartley strain; Charles River Laboratories, Inc., Wilmington, MA), each weighing approximately 350 g, were anesthetized with a combination of pentobarbital sodium (Nembutal; 25 mg/kg, intraperitoneally), fentanyl (0.2 mg/kg intramuscularly) and haloperidol (10 mg/kg, intramuscularly). Half doses of each drug were given at first. One third or a quarter of the full dose was administered as needed to maintain the appropriate depth of anesthesia. After completion of the experiment, animals were euthanized using a fatal dose of pentobarbital. A cannula for insertion into a cochleostomy was fabricated from PEEK, PTFE, and Tygon tubing. The infusion tip comprised a 3–cm length of 150–μm ID PEEK tubing connected to a 1.1–cm length of 101–μm ID PTFE tubing via a 4–mm length of 0.25–mm ID Tygon tubing. The PTFE end of the cannula was inserted into the scala tympani via a cochleostomy. The PEEK end of the cannula was connected to 0.25–mm ID Tygon tubing coming from the micropump. To prevent the PTFE tubing from inserting too deeply and to help seal the cochleostomy during infusion, a small droplet of silicone glue was placed 3 mm from the inserted end of the PTFE tubing and allowed to dry prior to insertion. We used a bullectomy approach and made a cochleostomy approximately 0.5 mm distal to the round window. The cannula coming from the micropump was inserted into the cochleostomy and glued to the bulla. Dental cement (GC Fuji 1, GC America, Alsip, IL) was used to seal the cochleostomy. A perfluoroalkoxy-alkane-insulated silver wire electrode (203 μm uncoated diameter, A-M systems, Carlsborg, WA) was placed near the round window niche and glued to the bulla. 2.6 Hearing tests CAPs and DPOAEs where measured under anesthesia at characteristic frequencies of 2.78, 4, 5.6, 8, 12, 16, 24, and 32 kHz. Acoustic stimuli were digitally generated and auditory responses were monitored using a National Instruments PXI stimulus generation/data acquisition system (National Instruments, Austin, TX), which included a 24–bit digital I–O board (PXI–4461) and a 16–bit digital I-O board (PXI–6221). Acoustic stimuli were delivered via a custom acoustic assembly comprising two dynamic drivers as sound sources and a miniature electret microphone to measure the sound pressure level (SPL) in situ. The system was controlled and measurements were recorded using the Eaton–Peabody Laboratories Cochlear Function Test Suite, a LabView-based software that is available online (Eaton-Peabody Laboratories 2014). CAPs were recorded in response to tone-pip stimuli (0.5– ms duration, 0.5–ms rise-fall; cos2 onset envelope; 16/s). The response from the electrodes was amplified (10,000x), filtered (0.3–3 kHz bandpass), and averaged (128 samples

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at each frequency–level combination). The SPL was incremented in 5 dB steps, from ∼10 dB below the estimated threshold to 100 dB SPL. DPOAEs were recorded for primary tones, f 1 and f 2, where f 2 was the characteristic frequency. The ratio of f 2 to f 1 was set to 1.2, and the SPL of f 2 was 10 dB less than the SPL of f 1. Primaries were incremented together in 5 dB steps from ∼10 dB below the estimated threshold to 80 dB SPL. The SPL in the ear-canal was measured, amplified, and digitally sampled at a rate of 200 kHz. 2.7 Acute drug infusion experiments For animal experiments involving acute infusion of DNQX into a guinea pig cochlea, the inlet of the micropump was connected to an external reservoir filled with DNQX solution (300 μM DNQX in artificial perilymph) via 10 cm of 0.25–mm ID Tygon tubing. The outlet of the pump was connected to the inlet of a flow sensor via 6 cm of 75–μm ID PEEK tubing (and short lengths of 0.25–mm ID Tygon tubing used as connectors). The outlet of the flow sensor was connected to a cannula (to be inserted in a guinea pig cochlea) via 100 cm of 0.25–mm ID Tygon tubing. The entire system was primed with the DNQX solution, where the pump was first primed by hand using a syringe, then the fluidic connection between the pump outlet and the rest of the system (PEEK, flow sensor, Tygon, and cannula) was established, and finally pump scheme 1 (fast infuse) was run until DNQX solution was dripping from the cannula. The pump was then shut off, and the cannula was inserted into a reservoir containing artificial perilymph solution. Pump scheme 4 (fast withdraw) was run until approximately 12 μL of solution were withdrawn (the flow sensor data acquisition software enables measurement of volume with an online integrator). This priming procedure resulted in AP comprising the 12 μL of fluid in the Tygon tubing closest to the cannula, which would be infused first. The rest of the fluid in the system was DNQX. After 12 μL of infusion, there would be a continuous transition from AP to DNQX. The cannula was then inserted into a cochleostomy made near the 28 kHz region of the cochlea, threaded apically 3 mm, and fixed in place with dental cement. DPOAE and CAP measurements were taken prior to the start of infusion to serve as a baseline. Pump scheme 3 (slow infuse – 5 min) was then initiated for 4 hours, and the flow rate was recorded. The target for pump scheme 3 was infusion of 1 μL over 10 s every 5 min, so artificial perilymph was expected to be infused for the first hour, and the DNQX solution for the remaining 3 hours. During infusion, we alternated between measurement of DPOAEs and CAPs, and repeated measurements as quickly as possible (one set of measurements was taken approximately every 30 min). Prior to each DPOAE measurement,

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the middle ear was checked for the build up of cerebrospinal fluid (CSF) leaking from the cochleostomy, and any fluid present was removed by aspiration. DPOAE and CAP measurements were conducted for up to one hour after the end of infusion, but were terminated earlier in cases where the guinea pig died under anesthesia prior to that point. 2.8 Data analysis The amplitude of a measured CAP waveform was calculated by subtracting the minimum of the waveform between t = 2 ms and t = 4.5 ms from the maximum of the waveform during the same time window. The restricted waveform was used in order to ensure that peaks in the waveform that resulted from frequency filtering in the system were not included in the amplitude analysis. The minimum input tone pip SPL that could generate a CAP amplitude of at least 0.4 μV was taken to be the CAP threshold of hearing (this value was empirically determined to be the the voltage at which CAP amplitudes began to rise linearly with increasing SPL in our system). If the CAP amplitude was less than 0.4 μV at the maximum tone-pip SPL of 100 dB, the threshold (and threshold shift) was set to 120 dB. Fourier analysis was used to determine the magnitudes of the DPOAE responses at f 1, f 2, and 2f 1 − f 2. In a DPOAE magnitude spectrum, for a given frequency, f , the noise floor was calculated as the average of the magnitudes at the 2 frequency samples above f and the two frequency samples below f . The minimum input SPL required to generate a 2f 1 − f 2 response with magnitude greater than the noise floor at 2f 1 − f 2 was taken to be the DPOAE threshold. DPOAE thresholds, CAP thresholds, and CAP amplitudes were grouped into 30–min bins so that they could be averaged across biological replicates. A single-factor ANOVA analysis was used to determine whether CAP amplitude suppression was frequency-dependent, and student’s t-tests were used to check for statistical differences between each of the measured frequencies.

Fig. 6 Volume delivered per stroke of the micropump as a function of load pressure applied at the outlet. Pump scheme 1 (fast infuse) was used to generate these data. DI water was the working fluid

maintained positive displacement up to a load pressure of 41 kPa (6 psi), but at 48 kPa (7 psi) the valves could not close and there was a large negative flow due to the pressure source. Instantaneous flow rates generated by the micropump during a pumping scheme were measured using a flow sensor. A typical flow rate profile for pump scheme 1 (Fig. 7), with 6 cm of 75–μm ID PEEK tubing connected inline as a fluidic resistor, shows that each actuator stroke was accompanied by a sharp increase in the flow rate. In this case, the average flow rate was 7.7 μL/min, and the peak flow rate was 9.3 μL/min.

3 Results 3.1 Pump performance characterization The valves in the micropump were sufficiently stable against a pressure of 20.7 kPa. The inlet valve leak rate was undetectable over 12 min of observation against 20.7 kPa, whereas the the outlet valve leaked at a rate of 262 μL/day against the same pressure. The micropump’s ability to generate flow against a pressure head was also characterized. Flow rates, and hence stroke volumes, were measured as a function of applied pressure at the outlet (Fig. 6). The pump

Fig. 7 Flow rate generated by the micropump as a function of time using pump scheme 1 (fast infuse) against an added hydraulic resistance of 1.29 kPa/(μL/min). 6 cm of 75–μm ID PEEK tubing was used to add the resistance, and the average stroke volume for the 7 strokes shown in these data was 0.082 μL. DI water was the working fluid

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Fig. 8 Total volume infused as a function of time during a demonstration of reciprocating flow. Pump scheme 2 (slow infuse – 1 min) was used to pump approximately 1.4 μL of DI water over 10 s, every minute for 5 min. Then pump scheme 4 (fast withdraw) was used to withdraw the fluid

Fig. 9 Flow rate generated by the micropump as a function of time using pump scheme 3 (slow infuse – 5 min) during infusion of AP/DNQX into a guinea pig cochlea. 6 cm of 75–μm ID PEEK tubing, 100 cm of 0.25–mm ID Tygon tubing, and the cannula were attached to the pump. One 10–s pump cycle is shown, comprising 8 actuator strokes. AP was the working solution

Flow sensor measurements were also used to demonstrate, as a proof-of-concept, reciprocating flow. For this demonstration, the outlet of the micropump was connected to a flow sensor via 6 cm of 75-μm ID PEEK tubing (a fluidic resistor). The entire system was primed with water, and then the inlet of the micropump was blocked. At this point, approximately 40 μL were stored within the fluidic capacitor (in total, 50 μL of water were in the pump). Pump scheme 2 (slow infuse – 1 min) was then initiated for 5 min while the flow rate was measured. This resulted in pumping of fluid stored in the capacitor out of the pump. Then the scheme was switched to pump scheme 4 (fast withdraw), which was run until all of the infused volume was withdrawn back into the fluidic capacitor. The flow rate measured during the reciprocation demonstration was integrated with respect to time to yield the volume infused as a function of time (Fig. 8). As intended, there was a step increase in the volume infused each minute for the first 5 minutes. After switching to the withdrawal pump scheme, all of the infused fluid was successfully withdrawn back into the fluidic capacitor, resulting in a net infusion volume of zero.

(Fig. 9). Measurements of instantaneous flow rates during infusion showed that in a typical flow profile for pump scheme 3, and the maximum flow rate did not exceed 10 μL/min. Integration of the flow rate with respect to time gives the cumulative infusion volume over the course of an acute infusion experiment (Fig. 10). Over the 4–hr duration of infusion, the micropump consistently infused slightly more than 1 μL every 5 min (1.1 ± 0.2 μL on average for 4 biological replicates). From these data, the

3.2 Pump performance during infusion into guinea pig cochleae Pump scheme 3 (slow infuse – 5 min), an infuse-only pumping mode, was used for all infusions into guinea pig cochleae. During these experiments, 6 cm of 75–μm ID PEEK tubing and a flow sensor were connected inline. Each ∼ 1 μL infusion comprised 8 pump actuator strokes

Fig. 10 Total volume infused as a function of time using pump scheme 3 (slow infuse – 5 min) during infusion of AP/DNQX into a guinea pig cochlea. 6 cm of 75–μm ID PEEK tubing, 100 cm of 0.25– mm ID Tygon tubing, and the cannula were attached to the pump. The inset shows a closeup of the first 15 min of infusion

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Fig. 11 (a) Example of a normal CAP for a guinea pig with good hearing, and (b) the (abnormal) CAP for the same guinea pig after 0.6 hours of infusion of 300 μM DNQX in AP using pump scheme 3 (slow – 5 min). The input tone pip was at 16 kHz and 90 dB for both waveforms

transition from AP infusion to DNQX infusion was estimated to occur at 0.87 ± 0.08 hours (average over 4 biological replicates), slightly earlier than the intended 1 hour.

We utilized DNQX, a glutamate receptor antagonist, as a test drug for intracochlear drug delivery via cocleostomy. We expected the DNQX to affect CAPs, but not DPOAEs (Chen et al. 2005). The DNQX served as a physiologic

indicator for drug delivery, as frequency-dependent changes in CAPs over time during DNQX infusion were used to infer information about the spreading of the drug along the length of the cochlea. AP was used as a control, in that no or minimal changes in CAPs or DPOAEs were expected during AP infusion. Infusion of DNQX decreased the amplitude of CAPs as expected (Fig. 11). The waveform shown in Fig. 11a is a typical CAP response to a tone pip, while the waveform shown in Fig. 11b is a typical response to a tone pip for which the CAP has been suppressed by the presence of DNQX. The largest two peaks in Fig. 11b are artifacts

Fig. 12 CAP amplitude as a function of time in response to 60-dB tone-pip stimuli during infusion of AP and DNQX into a guinea pig cochlea. Infusion began at t = 0, and amplitudes were normalized to the measurement taken just before t = 0. The gray shaded areas (t = 0 to t = 0.87 hours) represent the times during which artificial

perilymph was infused. The yellow shaded areas (t = 0.87 hours to t = 4 hours) represent the times during which DNQX was infused. The asterisks indicate data points that are the mean of either 3 or 4 biological replicates, and the error bars on those points represent the standard error of the mean. All other points have fewer than 3 replicates

3.3 Direct intracochlear delivery of DNQX (infuse only)

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Fig. 13 Time at which the CAP amplitude generated in response to a 60-dB tone pip dropped to 20 % of the initial value as a function of frequency. At t = 0 h, pump scheme 3 (slow infuse – 5 min) was initiated to start infusion of AP, and at t = 0.87 h, DNQX infusion began. The CAP amplitude did not drop to 20 % of the initial value at 2.78 kHz during the observational period for any replicate. Error bars indicate the standard error of the mean, and n = 4 biological replicates

from filtering in the acoustic system. We monitored CAP amplitudes as a function of time during infusion (Fig. 12) in order to observe the amplitude suppression effect of DNQX as it spread through the cochlea. Unexpectedly, CAP amplitudes began to decrease during AP infusion. The effect was the largest near the estimated cannula site (12–16 kHz), where CAP amplitudes decreased by a factor of 5 at the end of AP infusion. At all other frequencies, amplitudes decreased at the end of AP infusion by a factor of 2.5 or less. It is likely that the decrease in CAP amplitude was due to diffusion of DNQX into the AP when the system was primed, as a result of Taylor dispersion (see discussion). After the start of DNQX infusion, amplitudes just basal (32 and 24 kHz) and just apical (8 and 5.6 kHz) to the cannula location decreased by an additional factor of approximately 5. Toward the end of DNQX infusion, amplitudes in the cannula area (12–16 kHz) had decreased the most, whereas they decreased less at the most apical frequencies (4 and 2.78 kHz). After the pump was turned off, some recovery (an increase in amplitude) was observed at all frequencies except those near the cannula site (12 and 16 kHz). CAP amplitudes fell to less than 20 % of the initial value at frequencies near the cannula location (16 kHz) first, immediately after the start of DNQX infusion (Fig. 13). The effects quickly spread (within 30 min) to frequencies basal to that site. Frequencies apical to the cannula site experienced amplitude suppression later, where 80 % suppression was observed at 8 kHz after 45–60 min of infusion, and at 4–5.6 kHz only after 100 min of DNQX infusion. Amplitudes never fell below 20 % of baseline at the most apical frequency observed, 2.78 kHz. Analysis of frequencies from

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4–32 kHz with ANOVA confirmed that the duration of DNQX infusion required to reduce the CAP amplitude to 20 % of the initial value was frequency-dependent (p = 0.0004). Amplitudes did not fall below 20 % of the initial value during AP infusion. DPOAE and CAP hearing thresholds (Fig. 15) were also monitored as a function of time during infusion. Drug infusion did not have a significant effect on DPOAE spectra (Fig. 14). DPOAE thresholds did not change significantly over the course of the experiment, though there were some mild threshold elevations (∼ 25 dB) at all frequencies toward the end of the observation period, after the micropump was shut off. CAP threshold shifts generally followed trends similar to those observed for CAP amplitudes, but thresholds were less sensitive to DNQX infusion than amplitudes, as expected (Littman et al. 1989). During AP infusion, CAP thresholds did not increase significantly for most frequencies, but an increase of 25 dB was observed at 16 kHz just before the continuous transition from AP to DNQX infusion. Immediately after the start of DNQX infusion, the largest threshold increases were observed at 32 kHz, and increases were observed at all frequencies above and including 5.6 kHz. Thresholds did not increase significantly at 4 and 2.78 kHz, indicating that the dose of DNQX delivered to the apical region of the cochlea was not large enough to effect threshold changes, even though CAP amplitudes at those frequencies were reduced. After the pump was turned off, some recovery (reduction in threshold) was observed at all frequencies except those near the cannula site (12 and 16 kHz). Image plots of DPOAE threshold shift, CAP threshold shift, and normalized CAP amplitude (Fig. 16) provide visual representations of the data shown in Figs. 12 and 15, making it easier to see how the DNQX was distributed over time. There were no large changes in the DPOAEs (Fig. 16a) as a result of drug infusion, as indicated by the cool colors throughout the image. For the CAPs (amplitudes and thresholds), the drug effects were largest at frequencies basal to the cannula terminus (especially at 32 kHz) early in time, and the drug effect spread to apical frequencies (5.6–12 kHz) later in time, toward the end of DNQX infusion. After the pump was turned off, recovery was observed primarily at 24 and 32 kHz.

4 Discussion 4.1 Pump performance In guinea pigs with normal ears, the hydrostatic pressure in the perilymph ranges from 0.1–0.4 kPa (Yoshida and Lowry 1984; B¨ohmer and Andrews 1989; Andrews et al.

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1991; B¨ohmer 1991; Thalen et al. 2001), and there is a natural longitudinal flow of perilymph in the apical direction on the order of nl/min (Ohyama et al. 1988; Salt and Ma 2001). For successful intracochlear drug delivery, it is critical that the driving pump be capable of pumping against these pressures, and at rates that can produce significant drug concentrations at target locations given the natural clearance rate of the cochlea. We have demonstrated that our micropump generates positive flow on the order of 1 μL/min against a pressure of 40 kPa (Fig. 6). This leaves a significant margin for cases where clogs develop (potentially due to blood clots or biological particulate matter) leading to increased back pressure. There was essentially no decline in pump performance when our micropump was used to drive infusion of drug into guinea pig cochleae as compared to bench-top operation (Figs. 9 and 10). The pump was also very consistent in delivering precise injection volumes (Fig. 10). Across all infusion cycles in all of the biological replicates, the mean infusion volume was 1.1 ± 0.2 μL per cycle (n = 171 total boluses infused). One of the challenges with intracochlear drug delivery is the currently unknown potential for high flow rates and large infusion volumes to cause damage. Our group has estimated an upper limit on the peak flow rate for safe, acute drug infusion into a guinea pig cochlea in a pulsed delivery system (drug is delivered in short-duration bursts of flow at the peak flow rate that last on the order of hundreds of miliseconds, at intervals on the order of minutes) of 10 μL/min based on DPOAE data from previous work (Pararas et al. 2011). Our micropump is capable of generating flow rates beyond this safety limit, and in particular each actuator stroke results in a large spike in flow rate (Figs. 7 and 9). In the druginfusion experiments presented here, we chose to attach 6 cm of 75–μm PEEK tubing to the outlet of the micropump as a fluidic resistor; this resulted in peak flow rates that were Fig. 14 (a) Example of a normal DPOAE spectrum for a guinea pig with good hearing, and (b) the still normal spectrum for the same guinea pig after 1 hour of infusion of 300 μM DNQX in AP using pump scheme 3 (slow infuse – 5 min). f2 was at 16 kHz and 80 dB for both waveforms

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just below the estimated maximum limit for safety. If even lower flow rates are required, a larger fluidic resistance can be attached externally, or one can be incorporated into the pump’s design. The limited fluid space into which drugs can be delivered also presents challenges. In guinea pigs, the cochlear perilymph has a volume of about 8.7 μL and the scala tympani fluid space has a volume of about 4.7 μL (Thorne et al. 1999; Shinomori and Spack 2001). In the infusion experiments presented here, we infused relatively large volumes (40–50 μL) of fluid into the scala tympani over a 4–h period. While this fluid was expected to be cleared through the cochlear aqueduct and/or leakage from the cochleostomy site, it is currently unknown whether infusion of large volumes of fluid can lead to damage (no conclusive evidence of damage was observed in this study). Further investigation of this potential damage mechanism requires a reciprocating delivery system, like the one we have demonstrated as a proof-of-concept here (Fig. 8). If infusion of large fluid volumes does indeed lead to damage, such a reciprocating system will also enable zero-net-volume delivery, mitigating the effect. 4.2 Direct intracochlear delivery of DNQX (infuse only) DNQX, a well-known glutamate receptor antagonist, has been shown to reversibly disrupt hair cell to auditory nerve transmission when perfused into guinea pig cochleae (Littman et al. 1989; Sewell 1996; Parks 2000). Our results are consistent with previous reports (Chen et al. 2005; Pararas et al. 2011; Kim et al. 2014), where infusion of DNQX resulted in changes to CAPs (Fig. 11), but not to DPOAEs (Fig. 14). Partial recovery of CAP thresholds and amplitudes after the termination of DNQX infusion were also observed, which is also consistent with previous reports.

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Fig. 15 DPOAE (dashed lines, open circles) and CAP (solid lines, closed circles) threshold shifts as a function of time during infusion of AP and DNQX into a guinea pig cochlea. Infusion began at t = 0, and shifts were calculated with respect to measurements taken just before t = 0. The gray shaded areas (t = 0 to t = 0.87 hours) represent the times during which artificial perilymph was infused. The yellow shaded areas (t = 0.87 hours to t = 4 hours)

represent the times during which DNQX was infused. The asterisks (CAPs) and hash marks (DPOAEs) indicate data points that are the mean of either 3 or 4 biological replicates, and the error bars on those points represent the standard error of the mean. All other points have fewer than 3 replicates. In cases where there was no CAP response, CAP threshold shifts were set to 120 dB

The well-known tonotopic organization of the cochlea (Greenwood 1996) can be used to infer information about the distribution and spreading of drug (DNQX) in the cochlea based on the frequency-dependent CAP responses to stimuli. In our system, we infer that the cochleostomy was made in the 28 kHz region of the cochlea, which was near the cochlear aqueduct (Ghiz et al. 2001). Because we threaded the cannula 3 mm apically, we estimate that the

cannula terminated in the 12–16 kHz region. Based on the anatomy of the cochlea and the position of the cannula, we expected drug to flow basally from the cannula toward the aqueduct, resulting in fast distribution of the drug to frequencies higher than 12 kHz. Transport of drug to frequencies apical to the cannula terminus was expected to be primarily diffusive, and therefore significantly slower. The distribution and spreading of DNQX in our experiments as a

Fig. 16 Image plots of (a) DPOAE threshold shift, (b) CAP threshold shift, and (c) normalized CAP amplitude as functions of time and frequency. Warmer colors represent increased damage to hearing (larger threshold shifts, smaller CAP amplitudes). Pump scheme 3 (slow infuse – 5 min) was initiated at t = 0, and regions of preoperation (Pre), artificial perilymph infusion (AP), DNQX infusion

(DNQX), and no pumping (Off) are marked. For (c), the color bar is on a log scale, and the CAPs were generated in response to 60–dB tone-pip stimuli. Each pixel represents the mean of up to 4 biological replicates (Figs. 12 and 15 indicate the data points with more or fewer replicates)

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result of infusion, as inferred from our CAP measurements, were consistent with our expectations based on cochlear anatomy (Figs. 13, 15, and 16). Minor elevations (≤ 25 dB) in DPOAE thresholds were observed toward the end of infusion at frequencies below 8 kHz, and at all frequencies during recovery after the end of drug infusion. The cause of the low-frequency elevations is unclear; high flow rates and/or large infusion volumes are potential damage mechanisms (as discussed in Section 4.1) that could lead to such threshold increases. Longer-term chronic studies combined with reciprocating drug delivery are required in order to explore these putative damage mechanisms. The broad-spectrum elevations we observed during the recovery period are likely due to the animal nearing death from anesthesia during these long-term experiments. Only one animal survived long enough for DPOAE measurements to be taken after the termination of infusion. We observed slight decreases in CAP amplitudes immediately after the start of AP infusion, and moderate decreases toward the end of the AP infusion period. In previous reports (Barron et al. 1987; Bobbin and Ceasar 1987), CAP measurements after AP perfusion generally resulted in slight increases in CAP amplitude. The CAP amplitude suppression we observed during AP infusion was most likely due to DNQX diffusing into the AP, owing to the way in which we loaded AP into the system. In this study, we primed the entire infusion system with DNQX, and then withdrew 12 μL of AP into the cannula and Tygon tubing at an average flow rate 3 μL/min. During the withdraw process, Taylor dispersion resulted in a softening of the interface between the AP and the DNQX. Our estimates based on Taylor dispersion suggest that after the loading process, the 3 μL of AP nearest to the AP–DNQX interface had an average DNQX concentration of 45 μM. Additional dispersion that took place during infusion most likely caused distribution of DNQX to additional doses (we estimate that the last 5 1.1–μL AP doses had an average DNQX concentration of 42 μM). We can roughly estimate the effect that dispersion of DNQX into AP would have on CAP thresholds and amplitudes in our system by examining previous data regarding the dose-response relationship between DNQX and auditory-nerve CAPs. Previously, (Chen et al. 2005) characterized to some extent the dose-response relationship for DNQX in an infusion system that is similar to ours, and reported that a DNQX concentration as low as 100 μM was sufficient to cause CAP threshold elevations. From our Taylor dispersion estimates, only the last 1.1–μl dose of AP was likely to have had a DNQX concentration of at least 100 μM, though it is possible that lower concentrations of DNQX in our system could have led to CAP threshold elevations. Our data is consistent with only the later doses of AP having had significant concentrations of DNQX, as we

only observed CAP threshold elevations at the end of AP infusion at 12 and 16 kHz. CAP amplitudes, however, are more sensitive to DNQX than CAP thresholds. To our knowledge, there is no published data on the dose-response relationship between DNQX and CAP amplitudes in an infusion system that is similar to ours. However, data from a different perfusion system can provide insight regarding the sensitivity of amplitudes in comparison to the sensitivity of thresholds. Littman et al. (1989) reported a 50 % decrease in CAP amplitude after 10 min of perfusion with 8 μM DNQX, while a concentration of 33 μM was required to see threshold elevations. The greater sensitivity of CAP amplitudes vs. thresholds is also consistent with our data, as we observed CAP amplitude suppression during AP infusion, even though CAP threshold shifts were only observed at later times (along with greater amplitude suppression), when the concentration of DNQX was higher. Our proposed 5–actuator device (Fig. 2) will limit dispersion between drug and perilymph, as the drug reservoir will be isolated from the infuse-withdraw line with valves.

5 Conclusions We designed an infuse-withdraw micropump that is suitable for use as part of a system for long-term pharmacokinetic characterization of inner-ear drug delivery in guinea pig animal models. We fabricated and tested a simplified micropump and used it to demonstrate bidirectional reciprocating flow, on-board fluid storage, and drug infusion under load (including infusion into animal models). Our micropump maintained positive displacement of fluid against pressures much greater than those typically encountered during inner ear drug delivery via cochleostomy, making it robust against varying delivery conditions, including clogs that may form due to the collection of biomatter. We demonstrated safe, acute drug delivery via cochleostomy in guinea pig animal models by using our micropump to deliver the glutamate receptor antagonist, DNQX, and by monitoring the associated frequency-dependent elevations in CAP thresholds and suppression of CAP amplitudes. The safety of our pump for use in drug delivery was confirmed by the minimal changes in the DPOAEs during DNQX infusion. The micropump, along with an integrated drug reservoir, will be a key component in a long-term drug delivery system for animal models used for drug discovery. Such a system will enable highly controlled delivery of drugs and drug candidates over long periods of time while minimizing systemic exposure to the drug. The high degree of control offered by such a system will facilitate pharmacokinetic characterization of drug delivery, as metered boluses of drug

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are infused directly into the cochlea with this delivery system. Because many existing drugs and drug candidates are highly toxic when delivered systemically, localized, controlled delivery will be critical in treatment of auditory disorders in humans. Hair cell regeneration treatment will also likely require delivery of a timed series of reagents, so there will ultimately be a need for a miniaturized version of this type of microfluidics-based delivery system in an implantable/wearable format for humans. Acknowledgements The project described was supported by Award Number R01DC006848 from the National Institute on Deafness and Other Communication Disorders. The content is solely the responsibility of the authors and does not necessarily represent the official view of the National Institute on Deafness and other Communication Disorders or the National Institutes of Health. We would also like to thank Transon Nguyen for his assistance with laser machining.

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