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. Author manuscript; available in PMC: 2021 Jun 8.
Published in final edited form as: ACS Biomater Sci Eng. 2020 May 7;6(6):3335–3348. doi: 10.1021/acsbiomaterials.0c00065

In Vitro/Ex Vivo Investigation of Modified Utah Electrode Array to Selectively Sense and Pace the Sub-Surface Cardiac His Bundle

Ankur R Shah 1,2,§, Muhammad S Khan 2,§, Annie M Hirahara 1,2, Matthias Lange 2, Ravi Ranjan 1,2,3, Derek J Dosdall 1,2,3,4,*
PMCID: PMC7380681  NIHMSID: NIHMS1605624  PMID: 32715084

Abstract

Utah Electrode Arrays (UEAs) have previously been characterized and implanted for neural recordings and stimulation at relatively low current levels. This proof-of-concept study investigated the applicability of UEAs in sub-surface cardiac pacing, for the first time, particularly to selectively sense and pace the His-Bundle (HB). HB pacing produces synchronous ventricular depolarization and improved cardiac function. Modified UEAs with sputtered iridium oxide film (SIROF) tips (100 – 150 μm) were characterized for SIROF delamination using an electrochemical impedance spectroscopy (EIS), scanning electron microscopy (SEM), and voltage transient (VT) techniques at various current levels of up to 8 mA for a biphasic pulse with 1 ms duration per phase at 4 Hz. Our results indicate that at a short pacing duration of 20 s with current levels of up to 4 mA, the SIROF exhibited a strong charge-transfer performance. For the longer pacing duration (6 min), SIROF demonstrated its holding capacity at all current levels except for ≥2 mA when delamination commenced for the time exceeded 4 min (EIS) and 2 min (VT). UEAs were inserted in isolated, perfused goat hearts to record the HB electrograms in real-time. Both stimulated and unstimulated electrodes were characterized for SIROF delamination before, during and after in vivo work. Our findings indicate that UEA was stable during the heart’s contraction and relaxation phase. Further, at a short pacing duration with current levels of up to 4 mA, UEA demonstrated high selectively in sensing the HB. This proof-of-concept work demonstrates the potential applicability of UEAs in cardiac applications.

Keywords: Utah Electrode Array, Cardiac, His Bundle, Sensing, Electrochemical impedance spectroscopy, Scanning electron microscopy

INTRODUCTION

In recent years, His-bundle (HB) pacing has emerged as the leading candidate for physiological pacing because it provides nearly normal electrical activation of both ventricles and thereby avoids ventricular dyssynchrony (Kronborg and Nielsen, 2016; Sharma and Vijayaraman, 2016). HB pacing reduces the intraventricular and the atrioventricular dyssynchrony by providing a more physiological pattern of ventricular electrical activation (Kronborg and Nielsen, 2016; Lustgarten et al., 2015; Sharma et al., 2017). A single catheterized screw electrode is commonly used to access, sense and pace the HB (Gammage et al., 2006; Sharma et al., 2017). The location of HB in both the healthy and the diseased subjects is significantly varied which increases the difficulty in using a single electrode for clinical isolation of the HB (Kawashima and Sasaki, 2005). As the screw electrode is relatively large-sized, there is a potential for an injury to HB at the site of screw electrode insertion (Vijayaraman et al., 2016; Vijayaraman and Ellenbogen, 2018). Recently published and ongoing trials with HB pacing provide compelling evidence that HB pacing provides improved long-term outcomes for patients as compared to RV or biventricular pacing modalities with selective HB pacing achieved in approximately 1/3 of the successful implantations (Beer et al., 2019). It has been observed that a trained electrophysiologist must implant 30-50 devices before becoming proficient in successful HB pacing lead implantation (Keene et al., 2019). Thus, improved tools for lead implantation and selective HB pacing are needed to improve clinical uptake.

Utah Electrode Array (UEA) with sputtered iridium oxide film (SIROF) tips is one the most widely used implant for recording from and stimulating neurons for applications of the brain-machine interface and have demonstrated numerous both in vitro and in vivo applications at low current levels(Rajmohan Bhandari et al., 2010; Cody et al., 2018; Joshi-Imre et al., 2019; Kim et al., 2009; Straka et al., 2018; Vasudevan et al., 2017; Wark et al., 2013). The stimulation pulses used in cardiac applications are longer (~1 ms) (Sharma et al., 2017) as compared to neural pacing (~200 us). The stimulation current requirements are also much higher (0.1 – 4 mA) for cardiac applications in comparison with neural pacing (<100 μA) (Chugh, 2017; Claeys, 2014). The strength duration curve of an excitable tissue type provides the relationship between the pacing threshold and pulse duration. While the processes for fabricating smaller stimulating electrodes are well established, they are functionally limited by the amount of electrical charge that can be delivered through the smaller surface area.

The HB chronaxie is reported at 0.47 ms as compared to RV chronaxie of 0.79 ms in humans (Jastrzębski et al., 2019). Therefore, to use the UEA for a cardiac application, characterization needs to be demonstrated for biphasic pulses with 1 ms duration per phase separated by 0.2 ms. When implanted, the stimulation frequency for selective HB pacing to treat bradyarrhythmias is at or at most slightly faster than instrinsic sinus rhythm rate. The normal sinus rhtyhm rate dogs, goats, and rabbits is 100, 80, and 220 bpm (Flecknell, 1993). Since we currently consider the device for investigational use in animals hearts, 4Hz pacing was used in in vitro testing as the pacing rate is not expected to exceed 4Hz. Future adaptation of UEAs for cardiac applications will require detailed investigation to explore the impact of high current levels with longer stimulation pulses on the electrodes.

To investigate the stability of the UEA concerning SIROF delamination, we conducted electrochemical impedance spectroscopy (EIS) and voltage transient (VT) experiments in saline for in vitro characterization at higher current levels such as 1, 2, 4, 6 and 8 mA. For each subsequent current stimulation, modes were identified at higher current levels with short (20 s) and long duration (480 s) cycles. In vivo testing of the UEA was demonstrated in explanted perfused goat hearts to investigate the suitability of using UEAs in the HB-region for HB sensing and pacing. Scanning electron microscopy (SEM) was also used to capture images before and after electrochemical measurements in saline, and before and after in vivo work to study the SIROF stability and correlate it with the recorded impedance spectra. To the best of our knowledge, the proposed effort is the first proof-of-concept study in which UEAs were successfully characterized for biphasic pulses with 1 ms duration per phase separated by 0.2 ms at 4 Hz. Also, for the first time that the UEA was inserted in the HB region and HB electrograms recordings were demonstrated.

MATERIAL AND METHODS

Utah Electrode Array

Utah electrode arrays (UEAs) were custom designed and fabricated at Blackrock Microsystems, Salt Lake City, UT. Detailed fabrication steps have been described elsewhere (R. Bhandari et al., 2010; Rajmohan Bhandari et al., 2010). Briefly, the UEA consists of 100, 1.5 mm long needles arranged in a 10x10 matrix (Fig. S1a). The needles project out of a plaque that has a surface of 4 mmx4 mm. Each electrode needle is insulated along the length and also from the neighboring electrodes. Each electrode tip has a coating of SIROF with an estimated length ofas about 100 – 150 μm. The 96 electrodes of the UEA were interfaced with a Samtec connector through its bundle wiring (Fig. S1b). Optical microscopy was used for initial inspection of the fabricated UEAs. Detailed electrochemical impedance spectroscopy was employed at each stage of the experiments. Scanning electron microscopy was employed before and after experiments.

Electrochemical Impedance Spectroscopy

Electrochemical measurements were performed to characterize SIROF stability and recording capabilities over time before and after delivering higher current amplitudes through the electrodes. Both in vitro and in vivo electrode impedance measurements were recorded sequentially over the period under different stimulating current thresholds. All measurements were taken using a Gamry Interface 1010E (Gamry Instruments, Warminster, PA USA) in a standard three-electrode configuration (working, counter and reference). An individual electrode from the UEA was used as a working electrode and platinum wire (Millipore Sigma, St. Louis, MO) was used as a counter electrode. The impedance magnitude was directly accessed through the EIS recordings and analyzed at the signal voltage of 100 mV RMS amplitude spanning frequencies from 0.01 to 100 kHz.

Scanning Electron Microscopy

SEM data were captured using a Quanta 600 with a high-resolution field-emission source at an accelerating voltage of 10 kV. Electrodes were imaged under low vacuum in backscatter electron detector mode to visualize the SIROF stability and its delamination at different current thresholds. Images were taken before and after in-vitro electrochemical stimulation in saline, and before and after in-vivo HB sensing and pacing in explanted goat hearts.

Voltage Transients

The voltage transient measurements were performed in the saline at room temperature. To prevent DC leakage during the stimulation, an isolation capacitor (10 μF) was placed between the sourcemeter and working electrode. A voltage to current convertor (RdmApps, CO, USA) was used to generate a biphasic stimulating waveform of up to 10 mA at various current levels (0.1, 0.4, 1, 2, 4, 6 and 8 mA) with a 1 ms negative impulse followed by a 0.2 ms of a blank period and by another a 1 ms of positive impulse at 4 Hz rate. The current pulses were injected into the modified electrodes of the UEA and a digital oscilloscope (DSO2024A, Keysight, Santa Rosa, CA) was used to visualize and record the voltage transient with respect to reference. The maximum negative potential excursion was estimated to be the potential immediately after the end of the cathodic pulse. The maximum cathodic excursion (Emc), is equal to the electrode potential immediately after the end of the cathodic pulse (Negi et al., 2010). Similarly, the anodic electrochemical potential (Ema) is equal to the electrode potential immediately after the end of the anodic current pulse (Negi et al., 2010).

Animal preparation and surgery

All animal care and procedures were performed under the approval of the University of Utah Institutional Animal Care and Use Committee and in accordance with regulations specified by Guide for the Care and Use of Laboratory Animals (Albus, 2012). The goats (n=2) were anesthetized by intravenous injection of 6 mg/kg of Propofol (MWI Vet, Boise, Idaho). An endotracheal tube was used to intubate and provide a mixture of 2 – 3 % isoflurane and oxygen (at 2 L/min) for continued maintenance of anesthesia and respiration. The following parameters were monitored and maintained over the entire duration of the surgery: blood oxygenation level (spO2>95%), end-tidal carbon-dioxide level (etCO2 ~ 35), respiratory rate (12/minute), tidal volume (550 mL) and temperature (37.4 °C). After the monitored parameters stabilized, a medial sternotomy was performed to access and explant the heart from the animal. 5000AU of heparin was injected intravenously to prevent blood clot formation before excising the heart. The pulmonary veins, pulmonary arteries, superior and inferior vena-cava were cut, and the aorta was clamped before being severed. The heart was immediately submerged in a cold container of 4°C cardioplegic solution (which contained 110 mmol/L NaCl, 16 mmol/L KCl, 16 mmol/L MgCl2, 1.2 mmol/L CaCl2, and 10 mmol/L NaHCO3 ) until the aorta was cannulated and perfused.

Heart preparation and Langendorff perfusion

An aortic cannula was tied using umbilical tape in the aorta, and more cardioplegic solution was flushed through the heart while making sure no air bubbles were trapped in the pathway. All the extra tissues attached to the explanted heart were cleaned on benchtop. The mitral valve of the left ventricle was cut to prevent pressure buildup in the left ventricle. The heart was removed from the cardioplegic solution, and aortic connector was put in line with the Langendorff-setup which perfused the heart with Tyrode solution 4 mmol/L KCl, 1.8 mmol/L CaCl2, 130 mmol/L NaCl, 1.2 mmol/L NaH2PO4, 1 mmol/L MgCl2, 20.8 mmol/L NaHCO3, 11 mmol/L dextrose, and 0.04 g/L bovine albumin), (Sigma Aldrich, St. Louis, MO). The flow rate of the system was adjusted so that the inline pressure of the heart was maintained between 40-60 mmHg. The heart was also superfused with warm Tyrode solution, with temperature maintained at 37±0.5 °C. The Tyrode solution was aerated with O2 and CO2 to maintain a pH of 7.4±0.1. This preparation kept the explanted heart stable for at least 2-3 h depending on experimental protocols.

Instrumentation and experimental protocol

A pair of bipolar wire electrodes were inserted in the right ventricle to monitor the ventricular electrical activity. Another pair of bipolar wire electrode was inserted in the high left atrium to monitor the atrial activity. Three more electrodes were placed in the Langendorff setup’s bath so that they formed three corners of a triangle. These three electrodes were used to monitor the pseudo-ECG of the heart. The third electrode served as the common ground for the atrial, ventricular, and pseudo-ECG electrodes. The atrial and ventricular bipolar electrodes were used to verify that the atrium and ventricles beat sequentially and synchronously. The pseudo-ECG helped monitor the QRS complex of the ECG under different pacing conditions. The atrial and ventricular pseudo electrograms and pressure were being monitored using LabChart software through PowerLab 16/30 system (AD Instruments, CO Springs, CO, USA). Heart rate was estimated from the R-R interval of the pseudo-ECG trace.

UEA Insertion for HB Sensing and Pacing

An incision in the right atrium from the opening of the inferior vena cava towards the right atrial appendage exposed the basal right ventricular septum and the region containing the penetrating HB. A clinically used screw-electrode, SelectSecure 3830 (Medtronic, Minneapolis, MN), was inserted at the anterior vertex of Koch’s triangle to identify the location of HB (Fig. S2). The HB signal appeared on the electrode’s raw electrogram as a sharp deflection timed between the atrial activity and the ventricular activity. After identifying the location of HB, a pneumatic inserter (Blackrock Microsystem, Salt Lake City, UT) was used to insert the UEA into the location. The pneumatic inserter provided a punch on the backside of the UEA.

Integrated Instrumentation for HB sensing, pacing, and post-processing

The UEA was interfaced with a custom-designed electronic system that provides 92x preamplification for signal in the 0.03 – 5 kHz bandwidth. The system was used to acquire and visualize the data from 96 electrodes of the UEA in real-time at 4.096 kHz. Fig. S3 shows the two electrodes exhibiting real-time recording of sinus data with and without His signal sensing. After the data was acquired, spatial derivatives of all electrodes were visualized to see if the local HB signal stood out. Details of this method have been described elsewhere (Angel et al., 2014; Punske et al., 2003). The formula for the spatial derivative of an electrode is given by:

Vc−(Vw+Ve+Vn+Vs)4 (1)

Where V is the voltage, subscript ‘c’ is for the central electrode, and subscripts ‘w’, ‘e’, ‘n’, and ‘s’ are for electrodes west, east, north and south of the central electrode. Spatial derivatives highlight the electrical activity underneath the central electrode by subtracting the adjacent electrical signals and removing common far-field contributions to the recorded signal.

Any of the 96 electrodes can be selected on the system to pace through in a time precise manner. The current wave shape used for both in vitro and in vivo stimulation consisted of 1 ms negative impulse followed by 0.2 ms of a blank period, followed by a 1 ms of positive impulse. The time between stimulation pulses was determined by the frequency of pacing which may be fixed or determined from the activation rate at an electrode. The pseudo-ECG was used to distinguish the HB capture from the working myocardium (WM) capture by observing the shape of the QRS complex in the pseudo-ECG.

Statistical Analysis

A student’s T-test with unequal variances was conducted in PRISM 8.0 for all statistical testing. Results are written as mean±SD and expressed as significant (*p<0.05) and highly significant (**p<0.01).

RESULTS AND DISCUSSION

Electrochemical Impedance Spectroscopy

EIS experiments were conducted using a Gamry instrument in saline by sweeping the frequency from 0.01 to 100 kHz at a signal voltage of 100 mV. Based on the recently reported literature (Joshi-Imre et al., 2019), a hockey-stick region of the recorded impedance spectra was accessed from the raw data and was finally processed to investigate the SIROF stability and its delamination or any other potential electrode failure. The failures are attributable to a combination of factors, including SIROF delamination and damage or deterioration related to the wiring bundle. Wire bundle breakage includes potential failure sites at the connector board, along the length of cable, and where the bundle attaches to the UEA. Baseline impedance of all the 96 electrodes was recorded (6.43±1.48 kΩ at 1 kHz) in saline through EIS before testing them at different current levels. SEM was used (Fig. 1A and 1B) to capture both the side-view and the top-view of the UEA and the individual electrodes before in vitro testing.

Fig. 1.

Fig. 1.

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Characterization of UEA at higher current thresholds while stimulating for 20 s. (A) Cross-view and (B) Top-view of scanning electron microscopy images of UEA before electrochemical impedance testing and current stimulation. Insets show high-resolution SEM images of unstimulated electrodes. (C) Quantified impedance data at 1 kHz recorded through electrochemical impedance spectroscopy before (0 mA) and after (≥ 0.5 mA) stimulation at various current levels. Different current levels were used for stimulation to access SIROF stability and its delamination. Error bar in represents SD for n=9. ‘FR’ represents the failure rate threshold impedance where an electrode starts to delaminate its SIROF. (D) Representative SEM images of electrodes stimulated at various current levels for 20 s. SIROF started delaminating at 6 mA and 8 mA. *p<0.05 and **p<0.01.

A stimulating current of various levels utilizing a 1 ms negative impulse followed by a 0.2 ms of a blank period and by another a 1 ms of positive impulse at 4 Hz rate was employed to study SIROF stability and its delamination. For each continuous stimulation, the impedance spectrum was recorded subsequently. The recorded experimental data were quantified at 1 kHz to examine the failure modes of the 10 individual electrodes of the UEA as shown in Fig. 1C. The base impedance before stimulation for these 10 selected electrodes was 4.16±1.35 KΩ. After stimulation at 0.5 mA, there was a 37 – 50 % reduction in base impedance and was reported as 1.94±0.81 kΩ. The decreased in impedance after stimulating the stable electrodes is consistent with previously published findings (Otto et al., 2006). Stimulation was then continued at higher current levels of 1, 1.5, 2, 2.5, 4, 6 and 8 mA and the subsequent impedance was recorded as 1.94±0.81, 1.40±0.43, 1.28±0.31, 1.92±0.80, 4.33±3.61, 10.61±11.44, 20.51±14.36 and 26.59±16.56 kΩ at 1 kHz respectively.

If the electrode impedance increased by more than 2x after sending stimulation current pulses, the electrode was classified to have a weak SIROF holding capacity for the current level condition tested. If the impedance did not increase by more than 2x after sending stimulation current pulses, the electrode was classified to have a strong SIROF holding capacity for the current level condition tested. Failure rate (FR) was defined as rate at which the recorded impedance after a current stimulation surpassed at least twice the baseline impedance. These definitions were chosen carefully by inspecting the individual electrode tip morphology through SEM and its associated impedance in saline accessed through Gamry.

Interestingly, the electrodes showed strong SIROF stability at 0.5, 1, 1.5, 2, 2.5 and 4 mA, and weak SIROF holding capacity at 6 mA (p<0.05) and 8 mA (p<0.01) with pacing rate of 4 Hz for a continuous 80 pulses (20 s). The individual test had been repeated 9 times at each current level under similar conditions and the impedance spectrum was recorded after each stimulation. At higher current levels such as 6 mA and 8 mA, majority of the electrodes exhibited 5 – 7 times increase in impedance as compared to their baseline values. The Nyquist spectra of these electrodes exhibited variations in impedance response (Fig. S4) at lower frequencies (<100Hz) and were eliminated for further characterization and their utilization for sensing HB pacing during in vivo work.

To further explore the condition of the SIROF stability and its delamination at stimulated electrodes at various current thresholds, electrodes tips were studied under high-resolution SEM. Fig. 1D shows both the stable and delaminated response of SIROF at 1, 2, 4, 6 and 8 mA current levels. The increased impedance (>2x of the baseline impedance) and high-resolution SEM images demonstrated the effectiveness of SIROF both at lower and higher current pacing at 4 Hz. We also observed a wide range of SIROF delamination in another UEA’s electrodes at 6, 8 and 10 mA current levels under similar experimental conditions with a stimulating time period of 20 s (data not shown). Also, this UEA experienced cable and connector failures, to which we could attribute the abrupt loss of recording capability, and were not considered for further testing. The given data confirm that the UEAs have ability to be useful for in vivo HB pacing in large animal models of up to 4 mA pacing current for continuous 20 s but has limitations to pace the HB at ≥ 6 mA. We further tested the acceptable range of current thresholds (0.1 – 4 mA) for a longer continuous time period to examine if the UEAs could sustain SIROF integrity for longer-term HB pacing if required.

Fig. 2 shows the stimulation testing at different current levels of 0.1, 0.4, 0.7, 1, 2 and 4 mA for a longer duration. SIROF stability and its delamination were investigated in a similar fashion as recorded previously for 20 s. The impedance spectrum was recorded at the time interval of 10, 20, 30, 60, 120, 240, 360 and 480 s. The individual impedance was quantified at 1 kHz, as shown in Fig. 2 and Table S1, from the impedance spectra recorded at frequencies ranging from 0.01 to 100 kHz at a signal voltage of 100 mV. The UEA exhibited low impedance measures with 0.1, 0.4, 0.7 and 1 mA current levels for a longer period as shown in Fig. 2A, 2B, 2C and 2D respectively. However, for 2 and 4 mA, the SIROF started delaminating after 360 s, as shown in Fig. 2E and 2F respectively. These observations confirm that the developed UEA electrode tips remain stable when used for pacing at a current level of < 2 mA for at least 8 minutes. On the other hand, the UEA electrodes’ impedance increased by more than 2x the base impedance when they were paced for more than 360 s for the current levels of 2 mA (p<0.05) and 4 mA (p<0.01). These findings demonstrated that the UEAs originally developed for neuro applications have potential to be utilized in cardiac applications as an investigational tool, where sometimes a higher current threshold may be required temporarily to pace the HB.

Fig. 2.

Fig. 2.

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Electrochemical impedance outcome of UEA at 1 kHz over the period of time while stimulating at various current levels. (A) 0.1 mA. (B) 0.4 mA. (C) 0.7 mA. (D) 1 mA. (E) 2 mA. (F) 4 mA. SIROF showed strong holding capacity for 0.1, 0.4, 0.7 and 1 mA for continuous pacing of up to 8 min but it started delaminating at 360 s (6 min) while stimulating the electrodes at current thresholds of 2 and 4 mA. Error bars represent D for n=6 (A, D, E, F), and n=5 (B, C). Electrode numbers tested for each configuration listed in each plot. ‘FR’ represents the failure rate threshold impedance where an electrode starts to delaminate its SIROF. *p<0.05 and ** p<0.01.

Voltage Transients

To investigate the charge transfer performance characteristics (Emc and Ema) based on a 2 ms biphasic pulse at the 4 Hz rate with different current levels on the modified electrodes of the UEA, voltage transient analysis was conducted. Fig. 3 and Fig. 4 show the variation of Emc at different current levels over the period of time of various electrodes coated with SIROF. Previously, electrodes of the UEA have been characterized at 50 Hz with a pulse width of 200 μs and an interphase delay of 100 μs at the current levels of up to 20 μA (Negi et al., 2010), and the voltage transient parameters (Emc and Ema) were determined as the maximum negative and positive electrochemical potential limits associated with water reduction and oxidation, respectively. In another study (Park et al., 2018), the biphasic pulse of 100 μs duration followed by 100 μs inter-phase delay at the rate of 50 Hz. In both studies, they determined that the electrode potentials of Emc ≥ −0.6 V and Ema ≤ 0.8 V for the standard electrodes of the UEA were expected to cause dissolution or delamination of SIROF(Negi et al., 2010; Park et al., 2018). Based on their results, up to 0.6 V of Emc and 0.8 V of Ema were designated as the safe zone region with stimulating biphasic pulse of 200 μs, interphase delay of 100 s, 50-Hz frequency, with an incrementing amplitude of 0.5 or 1, up to 20 μA. In their work, they utilized the standard UEA with SIROF length on the tip of the electrode was about 50 – 70 μm.

Fig. 3.

Fig. 3.

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Voltage transient response data for Ema and Emc from six electrodes (n=4 each) having a constant current biphasic pulse for 20 s at different amplitudes (0.1, 0.4, 0.7, 1, 2 and 4 mA) at the rate of 4 Hz.

Fig. 4.

Fig. 4.

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Voltage transient data at different current levels over the period of time with a biphasic pulse width of 2 ms at 4 Hz. (A) 0.1 mA. (B) 0.4 mA. (C) 0.7 mA. (D) 1 mA. (E) 2 mA. (F) 4 mA. (Left) Voltage transient outcome of selected SIROF coated electrodes of the UEA in response to the biphasic pulse. A stimulating pulse was utilized at various current levels with a 1 ms negative pulse followed by a 0.2 ms of a blank period and by another 1 ms of positive pulse at 4 Hz. The maximum cathodic (center) and anodic (right) electrochemical potential (Emc and Ema) excursions of several electrodes during pulsing. The lowest Emc values which have shown a high potential in delaminating the electrodes were −0.876 V at 2 mA for t=240 s (E), and −0.915 V at 4 mA (F) for t=120 s. Similarly, the lowest Ema values which have depicted a high potential in damaging the electrodes were 1.244 V at 2 mA for t=240 s, and 1.281 V at 4 mA for t=120. For 0.1 mA (A), 0.4 mA (B), 0.7 mA (C) and 1 mA (D), the electrode potential (Emc and Ema) excursions were in higher than desired limit of −0.876 V and lower than 1.281 V respectively. Error bars represent data for n=4.

In our work, we characterized the modified UEAs with SIROF length on the tip was about 100 – 150 μm. A higher current threshold level of up to 4 mA was evaluated with a longer biphasic pulse of 2 ms at 4 Hz, and an interphase delay of 200 μs. Results showed that below 2 mA, electrodes exhibited a strong charge-transfer capacity in saline without damaging the SIROF as shown in Fig. 3. While pulsing with current less than 2mA for 20 s, Emc was higher than −0.6 V and Ema was lower than 0.8 V (Fig. 3). These results are in strong agreement with previously reported electrode potentials of Emc and Ema (Negi et al., 2010; Park et al., 2018). Stimulating at 2 mA and 4 mA, electrode potential (Emc and Ema) excursions were increased to (−0.746±0.244 and 0.92±0.141 V) and (−0.788±0.244 and 1.25±0.141 V), respectively. Since the effective surface area of SIROF was larger in the modified UEAs in comparison to the standard UEAs, Emc exhibited a higher cathodal-charge surface capacity. The similar phenomena were recently reported where the choice of the deposited material, electrode geometry, stimulating pulse width, and its duration highly influence the water window limit for both Emc and Ema (Cisnal et al., 2019; Park et al., 2019).

To further investigate the outcome of Emc and Ema for the longer duration stimulation, experiments were carried out for up to 8 min as shown in Fig. 4. The lowest Emc values which showed a high potential in delaminating the electrodes were −0.876 V at 2 mA for t=240 s (Fig. 4E), and −0.915 V at 4 mA (Fig. 4F) for t=120 s. Similarly, the lowest Ema values which demonstrated a high potential in damaging the electrodes were 1.244 V at 2 mA for t=240 s, and 1.281 V at 4 mA for t=120. For 0.1 mA (Fig. 4A), 0.4 mA (Fig. 4B), 0.7 mA (Fig. 4C) and 1 mA (Fig. 4D), the electrode potential (Emc and Ema) excursions were in higher than desired limit of −0.876 V and lower than 1.281 V, respectively. The summary of these results is shown in Fig. 4. The presented in vitro results suggest that that charge injection limits for any electrode material are highly dependent on the electrode bias level, waveform symmetry, current density employed, pulse frequency, and by properly choosing the stimulation protocol.

The results presented in this work show that Emc (> −0.876 V at 2 mA and > −0.915 V at 4 mA) and Ema (< 1.244 V at 2 mA and < 1.281 V at 4 mA) were not expected to cause dissolution or delamination of SIROF of the modified electrodes of the UEAs with the stimulation pulses used in these experiments, and the same has been observed during our electrochemical impedance recordings at 1 kHz for 2 mA and 4 mA (Fig. 2).

In vivo HB electrogram recordings through the UEA

The location of the HB is superior to the tricuspid valve, and by the anterior vertex of the Koch’s triangle. Selective HB pacing is defined as pacing that stimulates the HB directly without also capturing the local myocardium. Selective HB pacing can be confirmed if the following two conditions are met: (1) Shape of QRS complex on the pseudo-ECG traces from sinus beats and HB paced beats are very similar (2) The R-R interval becomes equal to the pacing interval. On the other hand, non-selective HBP is defined based on capture of basal ventricular septum in addition to HB capture. Non-selective HB is confirmed when : (1) no isoelectric interval observed between pacing stimulus and QRS complex on pseudo-ECG, (2) the shape of the QRS complex is broader than shape of spontaneous QRS complex on pseudo-ECG, and (3) The QRS complex narrows with higher-level current output (Sharma et al., 2017; Vijayaraman et al., 2016). The schematic in Fig 5A shows the relative location of the penetrating HB is below the AV node. The HB was made accessible by making an incision from near the coronary sinus ostium (CSO), as shown in inset of Fig. 5B, towards the right atrial appendage (RAA) and opening the right atrial flap. After ascertaining the location of the penetrating HB using a standard screw electrode, SelectSecure 3830 (Fig. S2), the UEA was inserted and stabilized at the HB location with the help of a pneumatic inserter as shown in Fig. 5B. Fig. 5C reveals the top view of the SIROF coated electrodes’ tips of the UEA before insertion.

Fig. 5.

Fig. 5.

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Conduction system and location of the His-Bundle (HB) region (A) HB region-of-interest is located below the AV node and is in the conduction path before the bundle branches. (B) UEA inserted in the RA above the tricuspid valve (TV) at the location where HB was initially identified using a screw lead (Fig. S2). The incision from the inferior vena cava opening near coronary sinus ostium (CSO) towards the right atrial appendage (RAA) opened the RA. Inset shows that the electrodes are intact within HB-region to sense and pace the HB. (C) SEM image of the UEA array shows the direction of insertion towards HB-region. Inset reveals high resolution (50 μm) morphology of SIROF before insertion and used for sensing and pacing the HB.

Fig. 6A shows the detailed electrograms recorded through 96 electrodes of the UEA implanted at HB-region. Two electrodes (# 10 and 58) showed either no or bad recordings and were excluded from the post-processing analysis. About 45% (43/96) of electrodes exhibited a successful HB signal deflection, as shown in Fig. 6B. The remainder of the electrodes showed normal electrograms comprised of only atrial and ventricular signals. Selected electrodes from the electrograms were used to stimulate the HB. In the second animal study, the same UEA was implanted using similar protocols, and HB signals were accessed. Fig. 7 demonstrates the sensing and pacing capability of the inserted UEA showing electrograms with a strong HB signal deflection (Fig. 7A) and the QRS (Fig. 7B) in the second explanted goat heart respectively. The HB signal is between the atrial (A) and ventricular (V) signals on the electrogram as shown in Fig. 7A (top). The center trace of Fig 7A shows the spatial derivative of the electrode. The spatial derivative removes the common electrical interferences between the central electrode and the surrounding electrodes highlighting only the local electrical activity of HB. Through the spatial derivative, the HB activity is most prominent and discernable, while atrial and ventricular activity still visible is less prominent than in the top trace. Fig 7A (bottom) shows the time derivative of the spatial derivative signal. In this trace, only the HB signal is prominent, while atrial and ventricular signals are not visible at all. During sinus rhythm, activation propagates through the conduction system to the working myocardium (WM). Therefore, the conduction system activation temporally precedes the activation of the local ventricular myocardium (Coronel et al., 2000).

Fig. 6.

Fig. 6.

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Raw electrograms recorded through UEA inserted in HB location. (A) 96 electrodes reveal Atrial (A), His (black arrows) and ventricle (V) signals. (B) 96 electrodes orientation and selected electrodes (yellow) with clear His signals.

Fig. 7.

Fig. 7.

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Electrograms recorded using an electrode of the UEA implanted at HB-region. (A, top) The Atrial, His-Bundle and Ventricular (V) activations are shown by the arrows (blue) and marked as A, HB and V respectively. (A, center) The spatial derivative highlighted the electrode’s local electrical activity by subtracting the surrounding electrode’s electrical signal from the central electrode’s electrical information. (A, bottom) The time derivative of the spatial derivative signal highlights the HB electrical activity (bottom). (B) pseudo-ECG while pacing through an electrode of the UEA. ‘P’ indicates the pacing artifact. T1=pace interval and T2=R-R interval. (B, top) Pacing HB at lower current utilizing an electrode from UEA, T1≠T2. (B, center) Pacing continued until HB captured, T1=T2. Inset shows that QRS complex while HB captured and is of similar morphology as the QRS of pseudo-ECG when HB is not captured. (B, bottom) UEA’s electrode is on working myocardium location beside HB-region. Instead of pacing HB, WM was paced and captured, T1=T2. The inset shows the QRS complex while WM is captured and the morphology is significantly different in comparison with QRS complex while HB is captured. QRS insets of HB and WM paced beats are of different shapes with durations of 32 ms and 40 ms, respectively. Electrograms recorded through 96 electrodes of UEA while implanted in HB-region are shown in Fig. 6.

Further, the time derivative of the spatial derivative simplifies the detection of local activation time and provides some immunity to noise inherent in these signals. Details of both spatial and time derivatives have been described in the methods section. This exercise helped in identifying the electrodes that were directly over the HB and could be useful for selective pacing.

After identifying the selected electrodes of the UEA directly on the HB, the electrodes of interest were paced at a rate slightly faster than the intrinsic heart rate (Chugh, 2017). Fig 7B shows pseudo-ECG traces under different pacing conditions. All the electrodes were paced at 2 Hz, and pacing artifacts are shown as ‘P’ in Fig 7B. Fig 7B (top) shows the pacing interval (T1=500 ms) and the R-R interval (T2=632 ms). HB location was initially paced at a lower current level of 0.2 mA and gradually increased at the rate of 0.2 mA/s until the R-R interval reached the pacing rate. At a current of 4 mA, the pacing interval and the intrinsic R-R interval are equal (T1=T2=500 ms) as shown in Fig. 7B (center). The inset shows the QRS morphology during selective HB capture and is similar while the HB was not captured (Fig. 7B, top). On the other hand, instead of pacing HB, Fig. 7B (bottom) shows the pacing traces while T1=T2=500 ms with WM captured. This region is located just adjacent to the HB (Fig. 5A) and this confirmed that the specific electrode is non-selective for HB pacing. Inset shows a QRS morphology for WM which is different in comparison to QRS morphology for HB. The QRS during HB pacing was 32 ms while the QRS during WM pacing was 40 ms long.

UEA stability while sensing and pacing HB

To investigate the electrodes tips’ SIROF stability and delamination while sensing and pacing the HB, only those electrodes were accessed that displayed a confirmed HB signal along with atrial and ventricular sensing as previously shown in the raw electrograms (Fig. 7). Prior to in vivo work, the impedance of all the electrodes of the new inserted UEA was recorded using a Gamry instrument as per protocols defined in the methods section. Data processing, analysis, and quantification have been performed in a similar fashion as described for the other UEA (Fig. 1 and Fig. 2) to evaluate the viability of these electrodes both at lower and higher current levels. Fig. S1b shows the mean base impedance magnitude (n=3) of the electrodes recorded in saline at 1 kHz before implantation. The impedance spectra were acquired over a frequency range of 0.01 to 100 kHz with a step size of 10 points per decade. Fig. 8 shows in vivo data collected from two explanted goat hearts to study the proof-of-concept feasibility of the UEA in sensing and pacing the HB at higher current thresholds. As described previously (Joshi-Imre et al., 2019), a hockey-stick region (0.01 to 100 kHz) of the impedance spectra was accessed from the raw data and further processed. Fig. 8A and Fig. 8C show the experimental in-vivo impedance outcome at 1 kHz for both unstimulated and stimulated electrodes of the UEA while sensing HB, respectively. Herein, the word ‘stimulated’ refers to those that were later used in pacing the HB. The in vivo impedance of ‘unstimulated’ (n=5) and ‘stimulated’ electrodes (n=5) while sensing HB before pacing was 32.23±10.18 kΩ and 30.26±8.81 kΩ respectively, as shown in Fig. 8E. Experimental values were measured directly through the Gamry instrument instead of utilizing any experimental fit based on equivalent circuit models such as Randles circuit to extract Tyrode solution resistance (Rs) or relevant capacitance details. Similarly, in vivo impedance was later recorded after pacing the HB with unstimulated (Fig. 8B) and stimulated (Fig. 8D) electrodes. The in vivo impedance of unstimulated (n=5) and stimulated electrodes (n=5) while sensing HB after pacing was 41.42±12.84 KΩ and 5.36±4.11 KΩ respectively, as shown in Fig. 8E. The electrodes, as shown in Fig. 8D, exhibited a lower impedance after current stimulation. Neither the impedance nor recording properties of the neighboring unstimulated electrodes had changed. Fig. 8F shows the relative change in the net impedance of unstimulated and stimulated electrodes after pacing the HB-region.

Fig. 8.

Fig. 8.

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In vivo electrical impedance of stimulated and unstimulated electrodes of the UEA embedded in the HB region of two explanted goat hearts. (A) Impedance spectra for unstimulated electrodes before HB pacing. (B) Impedance spectra for unstimulated electrodes after HB pacing. (C) Pre-impedance spectra of selected electrodes before HB pacing. Herein, these electrodes referred to those that were actually used in pacing HB and the change in impedance was reported in (D) as post-impedance spectra of stimulated electrodes after HB pacing. (E) Impedance data comparison accessed at 1 kHz for unstimulated and stimulated electrodes before and after HB pacing. Here orange and light-orange refer to in vivo pre- and post-impedance mean data of unstimulated electrodes, respectively, that have not participated in HB pacing. Blue and light-blue refer to in vivo pre- and post- impedance mean data of stimulated electrodes that were participated in HB pacing. (F) Relative change in impedance of unstimulated and stimulated electrodes sensing HB after pacing the HB-region. Substantial decrease in impedance was observed in stimulating electrodes utilized in pacing HB in comparison to nearby electrodes that were not stimulated (control). Error bars in (E, F) represent SD for n=5. ** p<0.01.

In comparison to pre-impedance data, stimulated electrodes showed 80.7±14.6 % relative change in impedance after pacing. On the other hand, unstimulated electrodes exhibited 29.7±20.1 % relative change in impedance data before and after pacing the HB. The present work also showed that the UEA was stable at the insertion site during the heart’s contraction and relaxation phases over the entire duration of the in vivo experiments while sensing and pacing the HB (Fig. S6).

In addition to recording in vivo impedance of selected stimulated and unstimulated electrodes inserted into the HB region, electrode surface properties were further characterized using high-resolution SEM imaging. Fig. 9 shows the high-resolution SEM images of selected electrodes of UEA that were used in HB sensing and pacing. Fig. 9a reveals the general morphology of the central region of the UEA before and after in vivo work was conducted in two explanted goat hearts. None of the electrodes were broken during insertion and removal of UEA from the heart and showed stability over the period while the HB was accessed. Unstimulated electrodes did not show any change in their morphology before and after in vivo testing. However, at higher current levels (Fig. 9b), electrodes morphology had changed due to the delamination of SIROF.

Fig. 9.

Fig. 9.

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High resolution SEM images of selected stimulated and unstimulated electrodes of UEA used in in vivo recording. (A) Randomly selected electrodes morphology before (left) and after (right) insertion in HB-region of isolated perfused goat hearts. (B) Selected stimulated electrodes that showed HB signals along with atrial and ventricular sensing before (top) and after (after) pacing the HB at 4 mA, 6 mA and 9 mA.

CONCLUSIONS

This proof-of-concept study demonstrated for the first time that a matrix of needle-like electrodes is useful to identify specific electrodes that are directly on the HB location in larger animal hearts such as goats. We extensively characterized the UEAs through EIS, SEM, and VT to investigate the SIROF stability at low and high current levels by pacing for biphasic pulses with 1 ms duration per phase separated by 0.2 ms. Our results demonstrated that the electrodes have strong SIROF holding capacity for 0.1, 0.4, 0.7 and 1 mA for continuous pacing of up to 8 min (validated through EIS, SEM, and VT). Whereas, at 2 and 4 mA, EIS exhibited that the electrodes started delaminating at 6 min and 4 min which were contrary to the VT results (4 min and 2 min, respectively). These differences could be due to the manufacturing variability of the electrodes investigated in individual experiments. For in vivo work, the UEA successfully held its position by the multitude of needles embedded in the tissue. The present work showed that the UEA was stable during the heart’s contraction and relaxation phase while selectively sensing and pacing the HB. The spatial configuration of the UEA help in identifying highly localized electrical activation times through spatial derivatives and their time derivatives. Both unstimulated and stimulated electrodes’ impedance and imaging were studied before and after pacing the HB. This proof-of-concept work recommends further validation of both EIS and VT results at the higher current levels of 2 and 4 mA for a longer duration, and adaptability of the UEA both in sensing and pacing applications involving the HB.

Supplementary Material

Electronic Supplement

Acknowledgments:

Research reported in this publication was supported by National Heart, Lung, and Blood Institute of the National Institutes of Health, NIH under award: R01HL128752 (DJD), a research grant from the Nora Eccles Treadwell Foundation (DJD) and American Heart Association, AHA under award: 9POST34450115 (MSK). The content is solely the responsibility of the authors and does not necessarily represent the official views of the NIH/AHA. The funders had no role in study design, data collection and analysis, decision to publish, or preparation of the manuscript. SEM was conducted at Surface Analysis Lab located at Utah NanoFab facility of the University of Utah, Salt Lake City UT.

Footnotes

Conflict of Interest: Authors declare no conflict of interest.

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