© 2017 The Royal Society of Chemistry

Organs-on-chips with integrated electrodes for trans-epithelial electrical resistance (TEER) measurements of human epithelial barrier function

By Olivier Y. F. Henry, Remi Villenave, Michael J. Cronce, William D. Leineweber, Maximilian A. Benz

and Donald E. Ingber

May 26, 2017Lab On A Chip

Lab on a Chip

PAPER

Cite this: DOI: 10.1039/c7lc00155j

Received 15th February 2017,Accepted 26th May 2017

DOI: 10.1039/c7lc00155j

rsc.li/loc

Organs-on-chips with integrated electrodes fortrans-epithelial electrical resistance (TEER)measurements of human epithelial barrierfunction†

Olivier Y. F. Henry, ab Remi Villenave,‡a Michael J. Cronce, a

William D. Leineweber, a Maximilian A. Benz a and Donald E. Ingber *abc

Trans-epithelial electrical resistance (TEER) is broadly used as an experimental readout and a quality control

assay for measuring the integrity of epithelial monolayers cultured under static conditions in vitro, however,

there is no standard methodology for its application to microfluidic organ-on-a-chip (organ chip) cultures.

Here, we describe a new microfluidic organ chip design that contains embedded electrodes, and we dem-

onstrate its utility for assessing formation and disruption of barrier function both within a human lung air-

way chip lined by a fully differentiated mucociliary human airway epithelium and in a human gut chip lined

by intestinal epithelial cells. These chips with integrated electrodes enable real-time, non-invasive monitor-

ing of TEER and can be applied to measure barrier function in virtually any type of cultured cell.

Introduction

Organs-on-chips (organ chips) are microfluidic cell culture de-vices that contain continuously perfused hollow micro-channels inhabited by living cells arranged to simulate tissue-and organ-level physiology that are currently being exploredas potential replacements for animal testing.1 One of themain advantages of this type of microphysiological culturesystem is that it enables high-resolution, real-time imaginganalysis of living human cell structure and function in atissue- and organ-level context in vitro.2,3 In organ chip stud-ies, one of the most useful measures of epithelial or endothe-lial tissue viability and function is the state of the permeabil-ity barrier. Measurement of trans-epithelial electricalresistance (TEER) is a quick, conventional, and non-invasiveassay that is used to evaluate the level of integrity and differ-entiation of in vitro epithelial monolayers in conventional

static cultures because the electrical impedance across an epi-thelium or endothelium is directly related to the formation ofrobust tight junctions between neighboring cells.4 While useof TEER measurements of cell cultures has become a stan-dard experimental method to estimate cell monolayer inte-grity, TEER measurements in organs on chips remains techni-cally challenging as no practical and validated approach hasbeen developed. This is because the closed, micrometer-sized,microfluidic channels that support cell growth and differenti-ation in organ chips restrict easy access to the epithelium andmake it difficult to carry out TEER measurements. Thus, it isvirtually impossible to record changes in permeability contin-uously using TEER in microfluidic culture systems.5

Patterned electrodes have been integrated into PDMS micro-fluidic devices in the past,6 but only a few have been used tomonitor epithelial barrier function in situ within organ chips.This was accomplished previously on-chip by direct insertionof metal wires into pre-molded locations above and belowmembrane-supported cell monolayers,7–12 repeated insertionof manually manipulated electrodes normally used to measureTEER in Transwell experiments,13 construction of cell culturechambers around large electrodes,14 or integrating glass orpolymeric substrates that contain electrodes formed using con-ventional metal patterning techniques into microfluidic culturedevices.15–17 As a result, the results of TEER studies with cellscultured in organ chips presented to date very often suffer fromlarge measurement variability, low sensitivity and they can behighly affected by non uniform cell cultures. Electrode locationalso can significantly alter TEER readouts in these cultures,

Lab ChipThis journal is © The Royal Society of Chemistry 2017

aWyss Institute for Biologically Inspired Engineering, Harvard University, CLSB5,

3 Blackfan Circle, Boston, MA 02115, USA.

E-mail: don.ingber@wyss.harvard.edu; Web: www.wyss.harvard.edu;

Fax: +617 432 7828; Tel: +617 432 7044bHarvard John A. Paulson School of Engineering and Applied Sciences,

Cambridge, MA 02139, USAc Vascular Biology Program and Department of Surgery, Boston Children's

Hospital and Harvard Medical School, Boston, MA 02115, USA

† Electronic supplementary information (ESI) available. See DOI: 10.1039/c7lc00155j‡ Current address: Emulate Inc., 27 Drydock Ave, Boston, MA 02210; www.emulatebio.com

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although mathematical models have been proposed that canhelp reduce these variations.5,17,18 Thus, successful fabricationof a robust on-chip TEER sensing capability would allowperforming simple electrochemical measurements to reliablystudy barrier function. In addition, because this type of methodmeasures electrical impedance, it also could be used to moni-tor ion channel activity, tissue conductivity, dissolved gases,cell proliferation, migration and many other cell behaviors. Inthe present study, we therefore set out to develop methods tofabricate organ chips with fully integrated electrodes.

To address this ongoing issue, we developed a new 2-chan-nel organ chip design with integrated electrodes that enablesreal-time measurement of TEER across virtually any type ofcell monolayer cultured inside the device. Here we show howto fabricate these TEER chips, and as a proof of concept, wedemonstrate that they can be used to continuously monitorthe formation and disruption of a well-differentiated humanlung airway epithelium growing inside a recently describedmicrofluidic human lung airway chip.19 We also confirmedthat similar TEER measurements can be carried out with hu-man intestinal epithelium cultured within the same chips.Thus, these TEER chips potentially can be used to non-invasively monitor the integrity of virtually any type of cul-tured tissue or cell monolayer in real-time. It also may pro-vide a convenient quality control assay to assess cell growthand differentiation inside microfluidic cell culture devices.

Material and methodsFabrication of TEER chips

To pattern the electrodes to fit within microfluidic organ chips,polycarbonate (PC) sheets (1 mm thick) were cut into 30 × 40mm substrates with their protective backing and inlets and out-lets drilled as required. After the protective backings were re-moved, the PC substrates were rinsed with isopropanol, dried ina stream of compressed air and activated in oxygen plasma for 2minutes (Technics Micro Stripper Series 220, 20 SCCM O2, 300mT, 100 W). The electrode patterns were laser cut in silicon-coated paper backing and consisted of two 1 mm wide electrodesseparated by 1 mm. The silicon side of the resulting papershadow masks were gently applied to the activated PC substratesusing a homemade alignment jig. These substrates were sequen-tially coated with 3 nm of titanium and 25 nm of gold in a metale-beam evaporator (Denton Vacuum LLC, USA). The papershadow masks were finally gently peeled off the PC substrates.

Polycarbonate and polyacrylic sheets were purchased fromMacmaster Carr (USA). PDMS Sylgard 184 was obtained fromDow Corning (USA). All chemicals were purchased fromSigma Aldrich (USA). Microfluidic channels were cut in 1 mmand 0.2 mm PDMS films prepared by spin coating ontoacrylic discs. The resulting PDMS coated discs were coveredwith low tack adhesive tape (Magic Tape 3M) to reduce sur-face contamination during subsequent processing and stor-age. The PDMS channels were defined using a CO2 laserusing minimal power to limit channel roughness and ashingof the thicker layer. Finally, the patterned PDMS layers were

cut to the size of the polycarbonate substrates using suffi-cient power to cut through both the PDMS layers and theacrylic disks, thereby producing a PDMS channel on anacrylic “stamp”. The protective tape was removed and chan-nels were cleaned with adhesive tape and isopropanol to re-move any loosely bound residue. The stamps were finally cov-ered with a new adhesive tape for storage.

The TEER Chip with integrated electrodes was then as-sembled following a layer-by-layer approach, schematicallypresented in Fig. 1A. PC substrates and PDMS stamps freedof their protective adhesive were activated in an oxygenplasma (20 SCCM O2, 300 mT, 100 W, 60 seconds) before be-ing immersed for 20 minutes in a 1% aqueous solution ofAPTES and a 1% aqueous solution of GLYMO respectively.PDMS and PC substrates were rinsed in water and dried in astream of compressed air before being aligned, brought incontact, and gently pressed together to ensure conforma-tional contact and baked at 60 °C overnight.

The acrylic backings were gently lifted off and the PC/PDMS assembly and plasma activated together with laser cutporous PET membranes (0.4 μm diameter pores). PC/PDMSand PET membranes were immersed for 20 minutes in a 1%aqueous solution of GLYMO and a 5% aqueous solution ofAPTES respectively, rinsed in water and dried in a stream ofcompressed air. The porous PET membrane was finallyaligned and brought in contact with a thin PC/PDMS (0.2mm) layer comprising the basal microfluidic compartmentand covered with a thicker PC/PDMS (1 mm) layer thatformed the apical microfluidic compartment. The assembledmicrofluidic chips were finally baked at 60 °C overnight(Fig. 1B). Porous PDMS membranes were prepared and as-sembled with PC/PDMS fluidic layers as previously de-scribed12 by sequentially exposing the various PDMS surfacesto oxygen plasma (30 seconds, 20 SCCM, 110 mT, 50 W) andbringing them into conformational contact. Finally, deviceswere cured at 60 °C overnight.

Epithelial cell culture

Our methods for culture and differentiation of primary humanairway epithelial cells (hAECs) in 2-channel microfluidic lungairway chips have been described previously.19 Briefly, hAECs(Epithelix; Switzerland) were expanded in one T75 cm2 tissueculture flask until ∼70% confluent using Lonza's BEGM growthmedium supplemented with growth factors (CC-3170, Lonza,USA). The porous PET membrane separating the upper andlower microchannels was coated with type I collagen (PureCol,Advanced Biomatrix, USA) and then hAECs were seeded (4.5 ×106 cells per mL) on its upper surface and incubated at 37 °Cunder static conditions to promote cell attachment. Mediumwas replaced daily with fresh medium for 6 days until the cellsbecome fully confluent and then an air–liquid interface (ALI)was generated to trigger mucociliary differentiation by remov-ing medium from the upper channel. The epithelium was fedby continuously perfusing (60 μL h−1) the lower channel withmedium for the duration of the experiment until full

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differentiation was attained (∼3 to 4 weeks). Healthy differenti-ated epithelium was maintained for 65 days, and the viabilityand quality of the epithelial cultures were assessed by monitor-ing epithelium morphology and integrity, cilia beating andpresence of mucus secretion by phase contrast microscopy, aspreviously described.19

In some studies, human Caco2 intestinal epithelial cellswere cultured within these TEER chips using previously de-scribed culture methods.12 In brief, the TEER devices were ex-posed to oxygen plasma for 30 seconds at a power of 50 Wusing a PE-100 plasma sterilizer (Plasma Etch, Inc. NV, USA),and then treated with 1% (3-aminopropyl)-trimethoxysilane(APTMS; Sigma) in 100% anhydrous ethanol for 10 min atroom temperature. After subsequent rinsing with 70% and100% ethanol washes, the devices were baked at 80 °C for 2hours. The insides of the channels were then coated with rattype I collagen (100 μg mL−1; Corning) in the presence ofMatrigel (10 μg mL−1) at 37 °C overnight in a 5% CO2 incuba-tor, followed by rinsing with DMEM culture medium. The in-testinal epithelial cells were seeded (1 × 106 cells per mL) onthe PDMS membrane and incubated at 37 °C for 3 days understatic conditions to promote cell attachment before initiatingstudies and maintained under flow (60 μL per hour) thereafter.

TEER measurements

A PGstat128N from Metrohm Autolab BV (The Netherlands)was used to record impedance spectra. Four-point impedancemeasurements were taken periodically over a period of 65days using a PGStat12/FRA (Autolab). Prior to initiating TEERmeasurements with cells cultured in organ chips, 50 μL ofwarm DMEM was gently introduced through the apical com-partment to wash away excess mucus. Following the washingstep, 50 μL of warm DMEM was introduced in the apicalcompartment and left to equilibrate at 37 °C for 10 minutes

before carrying out measurements. After measurement, theapical medium was again removed to restore ALI. TEER mea-surements were performed at room temperature for a maxi-mum of 2 minutes, which did not affect culture quality. Insome studies, EGTA (2 mM) was introduced in the apicalchannel, followed by the basal channel 10 minutes later todisrupt tight junctions, and chips were incubated at 37 °C for150 minutes. TEER values were measured every 10 minutesfor 1 h and then every 30 min thereafter at room tempera-ture. Cells were immediately placed back in incubator imme-diately after each measurement.

At high frequency (>10 kHz), the impedance curves aremostly characterized by the solution resistance, whereasTEER dominates the signal at lower frequency (<100 Hz); ca-pacitance is extracted from impedance data in the intermedi-ate range (100 Hz–10 kHz). This approach facilitates interpre-tation of TEER data as the background impedance of thesystem is automatically subtracted by the model. Severalmodels were tested and assessed based on both thegoodness-of-fit (χ2) criteria and their ability to provide a use-ful understanding of the underlying biology. The selectedmodel, which consists of a resistor RSOL in series with an-other resistor RTEER and a constant phase element (CPE),fitted all of the data well (χ2 < 0.01). More complex modelswith better χ2 were found to not always be able to fit all datasets or were difficult to interpret. CPEs are not typically usedin electrophysiology to model cell capacitance, but they havebeen shown to better fit the measured impedance of manycells.17 We also found this element to be particularly usefulto model the early stages of cell growth (<6 days). The mathe-matical expression of a CPE impedance is:

(1)

Fig. 1 A. Exploded CAD model of the TEER-chip. Gold electrodes are patterned onto polycarbonate substrates. Laser cut PDMS layers and PETmembrane are assembled using silane-based surface modification to irreversibly bond together. B. Photograph of the assembled TEER-chip, di-mensions are 25 × 40 mm.

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in which the CPE's impedance is expressed as a function ofthe system's admittance Yo, and an exponent n equaling 1 or0 for an ideal capacitor or an ideal resistor, respectively.Values for RTEER, Yo and n were estimated by modeling theexperimental data using the equivalent circuit presented inFig. 2B and using eqn (2) to calculate the capacitance of thecell layer Ccell expressed in farad (F).

(2)

Statistical analysis

All results are expressed as mean ± standard error. For thestatistical evaluation of quantified data, a paired t-test analy-sis was performed using GraphPad Prism version 7.0(GraphPad Software Inc., San Diego CA). Differences wereconsidered statistically significant when p < 0.05.

Results and discussionIntegrated TEER chip fabrication

We designed an integrated TEER chip (Fig. 1A and 2A) thatincludes four electrodes (2 above the upper channel and 2 be-low the lower channel) integrated into a microfluidic devicecontaining upper and lower channels separated by a porous(0.4 μm diameter) PET membrane or PDMS membrane, simi-lar to those that we previously used to create a human lungsmall airway chip19 and human gut chip,12 respectively. Theelectrodes were patterned on polycarbonate (PC) substratesusing a laser patterned, silicon-coated, backing paper,shadow mask. The electrodes were 1 mm wide spaced by 1mm, but well defined patterns as small as 0.3 mm can easilybe achieved using this simple technique. Shadow masks werecarefully peeled off their substrates and put in contact withplasma-activated PC sheets. A thin 3 nm titanium adhesivelayer was first evaporated onto the masked PC, followed by25 nm of gold. The resulting pattern was very stable to iso-propanol or ethanol rinse, as well as sonication for 5 minutesin buffer, and it was even difficult to remove using fingerpressure. Importantly, in addition to their stability, theseelectrodes proved to be transparent which allowed for quali-tatively assessing cell cultures using optical microscopy. In-lets and outlets were drilled on the top parts where required.

We chose PC as a base substrate for its high optical clarity,cell culture biocompatibility, ease of machining, compatibilitywith metal deposition processes and ease of chemical surfacemodification via silane chemistry to promote bonding to otherpolymer surfaces. Aminopropyl triethoxysilane (APTES) can read-ily interact at the surface of plasma activated PC through the for-mation of stable amide bonds between the acid groups generatedduring plasma treatment and the amine functionality ofAPTES.20,21 While bonding of plasma activated PDMS to thesilanol groups introduced at the PC surface via APTES treatmenthas been reported,22 we found that prior introduction of epoxymoieties at the PDMS surface using GLYMO resulted in improvedbonding23–25 and long term resistance to hydrolytic cleavage. Theporous PET membrane was then bonded to the open face of thePDMS layers using the same strategy, first modifying the PET sur-face with APTES and the PDMS with GLYMO before assemblingthe final device (Fig. 1B) and overnight curing at 60 °C. Thisresulted in a very stable bonding. As previously reported,24,25 theAPTES/GLYMO bonding can occur very rapidly at room tempera-ture, however we found that overnight at 60 °C produce more ro-bust devices. The difference in thermal expansion coefficient be-tween PDMS and PET did not result in the deformation of themembrane and it was found beneficial to flattening the mem-brane in the assembled device. The porous PDMS membranewas bonded to the PC/PDMS fluidic layers using oxygen plasmaand did not require any additional treatment.

Importantly, while glass also may be used to create micro-fluidic cell culture devices with integrated electrodes, PC hasmajor advantages in terms of its ease of machining. For exam-ple, electrodes can be easily patterned on PC or PET using tech-niques, such as conventional lithography and metal patterning

Fig. 2 A. Schematic view of the TEER-chip and 4-point impedancemeasurement chosen to measure TEER and capacitance. A small cur-rent of 10 μA of varying frequency is applied between two electrodes(Iexcite) located on each side of the cell culture, and the drop in poten-tial between the second set of electrodes measured (Vmeas). B. Exampleof impedance spectra recorded before ALI and after full differentiationat ALI. Lines are fitted data based on the equivalent electronic circuitused for modelling.

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techniques, as well as roll-to-roll laser ablation, which can con-siderably lower the cost of the final devices. Drilling and cut-ting glass also requires instruments that are not always avail-able in research laboratories. Another advantage of ourtechnique is that our chips can be assembled without requiringadditional layers for bonding, or to provide fluidic access.

TEER measurements in organ chips

Before initiating studies with organ chips, we carried out con-trol experiments measuring impedance of aqueous solutionsof 1, 10, 100 mM NaCl as well as the DMEM culture mediumwe use for our organ chip experiments (which contains 109mM NaCl, as well as various other components). These stud-ies confirmed that this impedance sensor is sensitive tochanges in conductivity (Fig. S1†). As expected, the imped-ance response recorded for DMEM was slightly lower thanthat of 100 mM NaCl as it holds a slightly higher conductiv-ity. TEER was then measured in 4 different organ chips, eachcontaining a well differentiated, mucociliated, human smallairway epithelium cultured in the upper channel at an air liq-uid interface (ALI) while the cells were fed by flowing (60 μLh−1) growth medium through the lower channel, as describedpreviously.19 All 4 lung airway chips were maintained for 62days in culture and 56 days at ALI, without any evidence ofcell toxicity from the presence of the gold and titaniumlayers. TEER measurements were carried out using a 4-pointsimpedance measurement (Fig. 2A) over the frequency range100 kHz to 10 Hz and the impedance curves fitted to anequivalent electric model (Fig. 2B). This simple model allowsextracting both TEER and cell capacitance values very effec-tively by factoring in the conductivity of the solution in everymeasurement (see methods for details).

TEER measurements were recorded at days 1, 4, 6, 17, 22,46 and 62 post seeding, and impedance and capacitancevalues were determined for each time point (Fig. 3). Analysisof these data revealed that during the first 6 days when cellswere cultured submerged in medium (before ALI), TEERvalues oscillated between an average of 200 Ohms (Ω) (day 4)

and 500 Ω (day 6), consistent with the progressive establish-ment of a tight monolayer of submerged airway epithelialcells. Creation of an ALI by removing the apical growth me-dium and flowing differentiation medium in the bottomcompartment resulted in a steady increase of TEER from 500Ω at day 6 (0 day at ALI) up to an average 1700 Ω at day 62(56 days at ALI). The capacitance of the cell layer followed asimilar trend, stabilizing after day 22 to an average value of220 nF. In addition, the value of n (see eqn (1) in methods)extracted from modelling the impedance response to theequivalent circuit described in Fig. 2B stayed under 0.5 overthe course of the 6 first days of culture (i.e., until the culturedcells reached confluence). Thereafter, n remained greaterthan 0.9, which can be associated with the formation of ma-ture cell monolayer, behaving more closely to an ideal capaci-tor. At those time points, substituting the CPE for a capacitorin the model did not result in significant changes to thevalue RTEER, RSOL and χ2. Importantly, the 10 μA excitationcurrent used for these measurements did not produce any de-tectable changes in cell behavior over 2 months of culture.

It is important to note that we present results here as Ω,rather than Ω cm2 (which normalizes TEER as a resistance de-pendent on the surface area of the cultured tissue being ana-lyzed), because while our TEER chip measures impedance ofthe culture area directly above and between the electrodes, italso is likely influenced by indeterminate regions of the cellculture outside the edges of the electrodes. While past TEERstudies on Transwell inserts and organs-on-chips normalizedtheir results for total tissue culture area and presented them asΩ cm2, most of those studies used hand-manipulatedelectrodes that can provide misleading results based on theirlocation or the stability of their position. There also was no evi-dence that the field was homogeneous across the entire tissueculture surface area in those studies, and so the biological sig-nificance of the Ω cm2 results they presented remain unclear.Thus, while presenting results in Ω might seem to be a limita-tion, the major advantage of our method is that the resultsobtained from one chip are directly comparable to all others

Fig. 3 Graph showing the TEER and capacitance data recorded overthe course of differentiation of the human primary airway epithelialcells. ALI = time of initiation of the air liquid interface (n = 4 individualchips).

Fig. 4 Impedance data for human Caco2 intestinal epithelial cellscultured in our TEER chip and recorded in the frequency range 0.1 Hzto 100 kHz.

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obtained with the same TEER chip design because of the ro-bustness of the electrode set-up. However, if investigators whouse these TEER chips in the future want to compare their re-sults with those obtained in past studies, they can estimate thevalue in Ω cm2 by multiplying the results they obtain by 0.03cm2 (i.e., the culture surface area above and between theelectrodes in our TEER chips), as long as they understand thecaveat that the actual surface area may be slightly larger.

To explore whether our TEER chip design can be general-ized to other cell types, we cultured human intestinal epithe-lium for 12 days within the same device, except that itcontained a porous PDMS membrane similar to that used ina previously described human gut chip.12 In these studies, theimpedance reached a stable plateau in the range 0.5 Hz to 100Hz, although some instabilities were observed at lower frequen-

cies, below 0.5 Hz (Fig. 4). Thus, these results show that theelectrical double layer does not interfere with our measure-ments in the frequency range studied, and they confirm thatthe method can be used with other types of epithelium.

To explore the dynamic measurement capabilities of theTEER chip, we then measured impedance and capacitancewithin the intestinal epithelium in the TEER chip beginning2 hours after seeding the cells and every day thereafter for 12days in culture (Fig. 5A). These studies revealed that impedancevalues rapidly increased to a plateau of 4046 ± 210 Ω by cultureday 3, while the capacitance stabilized at 182 ± 17 nF, whichcorresponds to the rapid formation of a stable intestinal epithe-lial monolayer. The chips were then placed under flow begin-ning on day 4, which initiated further differentiation and pro-moted the formation of three dimensional intestinal villi-like

Fig. 5 A. Overlay of impedance and capacitance values recorded during growth of intestinal epithelial cells and formation of an epithelialmonolayer on-chip (n = 5 chips). B. Plot of impedance values corrected based on measures of cell capacitance from A, which provide an estimateof total cell membrane surface area. C. Overlay of impedance recorded during exposure of the TEER chip lined by intestinal epithelium to 5 mMEGTA (black circles; n = 3 chips/condition) at different time points after treatment (the 24 h recovery period is highlighted in gray). TEER values ofcontrol chips cultured in parallel without EGTA were kept was determined at the same time points (open circles; n = 2 chips). All TEER values mea-sured for epithelium exposed to EGTA were significantly lower than control values (p < 0.0007) at all time points except for the initial and finaltime points. D. Phase contrast images of differentiated human intestine epithelial cells treated without (top) or with (bottom) 5 mM EGTA (bar,100 μm).

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structures, as previously described.12 Capacitance increasedvery rapidly thereafter to plateau at 1033 ± 269 nF by day 11,and this was associated with an increase in tissue surface areaduring differentiation due to villi formation, while impedancedecreased steadily to stabilize to 867 ± 131 Ω (Fig. 5A).

Capacitance correlates very well with cell membrane sur-face area and correlation factors have been proposed;26,27

thus, capacitance potentially can be used to correct imped-ance values for total cell membrane area. When we normal-ized impedance based on the measured capacitance values,we found that the corrected impedance values increasedsteadily during the first 4 days of culture, and then remainedat a steady plateau value thereafter (Fig. 5B).

To verify that the increases in TEER values we measuredwere specifically related to the presence of tight junctions, wetreated the intestinal epithelium both apically and basally atday 13 with or without EGTA (5 mM) for 2 hours duringwhich the impedance of the cultures was monitored regularly(Fig. 5C). Because EGTA is a strong Ca2+ chelator, it will affectadherens junctions as well as tight junctions, although defi-ciency in Ca2+ alone is not normally sufficient to induce theseeffects. When exposed to EGTA, the impedance of the epithe-lium dropped by 50% within the first 20 minutes andreached a maximum reduction of 65% by 1 hour (Fig. 5C),which is consistent with previous reports using conventionalTEER measurements in static cultures.28 In contrast, controlchips maintained near constant impedance values over thesame time course (Fig. 5C). EGTA treatment also was accom-panied by a loss of distinct boundaries around the villus-likeepithelial structures when contrasted with untreated gutchips (Fig. 5D), which is consistent with disruption of tightjunctions and loss of epithelial organization. Importantly, bywashing out the EGTA after 2 hours, and culturing the intes-tinal epithelium in control medium overnight, we could con-firm the reversibility of this effect by measuring impedanceusing the TEER chip during the recovery period and demon-strating restoration of impedance values by the end of theculture period that were indistinguishable from those mea-sured at the start of the experiment (Fig. 5C). Restoration ofnormal TEER values also was accompanied by reappearanceof more well defined villus-like structures, when analyzed byphase contrast microscopy (Fig. 5D). Control experiments car-ried out with EGTA and DMEM medium alone without cellsalso confirmed that EGTA has only a minimal direct effect(5.36 Ω increase in impedance at 12 Hz) on background im-pedance measurements (Fig. S2†).

Similar effects were also observed in human pulmonaryepithelium cultured in the lung airway chip 65 days post ALI.EGTA treatment (2 mM) resulted in a rapid drop in imped-ance values averaging 1800 Ω to 500 Ω within 150 minutes af-ter which impedance levels plateaued (Fig. S3A and B†). Thepresence of EGTA in the apical channel resulted in a very lim-ited drop in TEER believed to be the consequence of slowEGTA diffusion through the mucus layer,19 whereas subse-quent introduction of EGTA via the basal channel resulted ina more severe effect on the cell–cell junctions. These observa-

tions are in agreement with the mode of action of EGTA,which disturbs epithelial cell–cell adhesions, but does notsignificantly affect epithelial-matrix attachment or cellshape.29,30 This also was supported by phase contrast micro-scopic analysis that revealed EGTA treatment results in open-ing of the junctions between neighboring epithelial cellswithout causing obvious cell detachment, rounding or mono-layer disruption (Fig. S3C and D†).

Conclusion

The full integration of electrodes within organ chip micro-fluidic culture devices remains challenging. We developed asimple layer-by-layer fabrication process that enables assem-bly of complex microfluidic organ chips with integrated,semi-transparent, sensing electrodes to measure both TEERand cell layer capacitance using 4-points impedance measure-ments at varying frequencies. Using this approach, we wereable to follow the differentiation of human primary small air-way epithelial cells under ALI culture conditions and humanintestinal epithelial cells covered by flowing medium on-chip,as well as the disruption of cell–cell junctions and accompa-nying drop in TEER levels upon exposure to the chelatingagent EGTA. While we feel our system provided sufficientsensitivity, the location, dimensions and design of theelectrodes can be further modified to optimize excitation ofthe cell culture area and measurement of the electrical poten-tial across the tissue barriers depending on the organ chipdesigns utilized for culture. Nonetheless, our newly devel-oped TEER chip platform enables real-time measurements ofbarrier function that can be used to assess organ chip viabil-ity and changes in function in response to basal culture con-ditions, as well as drugs, toxins, inflammatory mediators, orother relevant external stimuli. This TEER chip technologyalso may open up new applications, including measurementof short circuit current, or action potentials of electrically ac-tive cells for electrophysiology studies, which are currentlylacking in the majority of microphysiological systems, thusexpanding the breadth of use of organ chips.

Competing financial interests

D. E. I. is a founder and holds equity in Emulate Inc., and hechairs its scientific advisory board. R. V. is currently an em-ployee of Emulate Inc.

Acknowledgements

This work was conducted with support from DARPA(#W911NF-12-2-0036), NIH (#5R01EB020004-03), FDA(#HHSF223201310079C) and the Wyss Institute for Biologi-cally Inspired Engineering at Harvard University. This workwas performed in part at the Center for Nanoscale System(CNS), a member of the National Nanotechnology Coordi-nated Infrastructure Network (NNCI), which is supported bythe National Science Foundation under NSF award no.

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Lab Chip This journal is © The Royal Society of Chemistry 2017

1541959. CNS is part of Harvard University. We thank AndresRodriguez for his assistance with graphic design.

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