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Published in final edited form as: Sens Actuators A Phys. 2021 Jun 1;331:112874. doi: 10.1016/j.sna.2021.112874

A micromachined force sensing apparatus and method for human engineered cardiac tissue and induced pluripotent stem cell characterization

Irene C Turnbull 1,*, Weibin Zhu 1, Francesca Stillitano 1, Chen-Chi Chien 1, Angelo Gaitas 1,*
PMCID: PMC8294102  NIHMSID: NIHMS1718827  PMID: 34305317

Abstract

Induced pluripotent stem cell derived-cardiomyocytes (iPSC-CMs) have great potential for cell therapy, drug assessment, and for understanding the pathophysiology and genetic underpinnings of cardiac diseases. Contraction forces are one of the most important characteristics of cardiac function and are predictors of healthy and diseased states. Cantilever techniques, such as atomic force microscopy, measure the vertical force of a single cell, while systems designed to more closely resemble the physical heart function, such as engineered cardiac tissue held by end-posts, measure the axial force. One important question is how do these two force measurements correlate? By establishing a correlation of the axial and vertical force, we will be one step closer in being able to use single cell iPSC-CMs as models. A novel micromachined sensor for measuring force contractions of engineered tissue has been developed. Using this novel sensor, a correlation between axial force and vertical force is experimentally established. This finding supports the use of vertical measurements as an alternative to tissue measurements.

Keywords: induced pluripotent stem cell derived cardiomyocytes, human engineered cardiac tissue, force sensing, micro-electromechanical systems (MEMS), axial force, vertical force

Graphical Abstract

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1. INTRODUCTION

Human induced pluripotent stem cells (iPSCs) are pluripotent stem cells reprogrammed from adult cells (e.g. adult human skin cells) which can then be differentiated into other cell types, such as cardiomyocytes (CMs) [1]. iPSC-derived cardiomyocytes (iPSC-CMs) have great potential for cell therapy [2, 3], drug assessment [4], and for understanding the pathophysiology and genetic underpinnings of cardiac diseases [5]. iPSC-CMs are emerging as a promising platform for drug evaluation [6] and hold significant potential for use in predicting cardiotoxicity [7, 8] and may find uses in personalized medicine [9]. Of particular importance is the use of iPSC-CMs as a human cellular model for predicting drug arrhythmogenicity [10]. Nearly one third of all drugs have cardiovascular side effects [11, 12].

Contraction forces are one of the most important characteristics of cardiac function and are predictors of healthy and diseased states [7, 13]. Several techniques have been used to measure the force of contraction from cardiomyocytes [7, 14–23]. However, most of these techniques are not useful in characterizing iPSC-CMs due to their irregular shape and smaller size; also, many of these techniques are unable to directly measure the force. An answer to this problem is the use of microcantilevers, such as atomic force microscopy (AFM) cantilevers [24, 25] and micro-cantilevers with embedded sensors [26, 27], that have already been used for cardiomyocytes contractile force measurements. These devices are like micron size diving boards that are brought in contact with single cells. Either using a laser light that is reflected off the cantilever, or by using an embedded sensor, the movement of the cantilever that corresponds to the cellular contractions is tracked. By characterizing the properties of the microcantilever it is a simple task to derive the force.

The cantilever techniques measure the vertical force of a single cell, while systems designed to more closely resemble the physical heart function, such as cardiac tissue on posts [28, 29], measure the axial force (Fig. 1). One important question is how do these two force measurements correlate? By establishing a correlation of the axial and vertical force, we will be one step closer in being able to use single cell iPSC-CM instead of more elaborate human engineered tissue or animal heart tissue as models. This would enable the use of single iPSC-CM for cardiotoxicity and personalized medicine. In this work, we attempt to establish such a correlation on tissue through measurements of human engineered tissue axial force using an established method [30] and vertical force using a micromachined cantilever with an embedded strain sensor. These simultaneous measurements of axial and vertical forces are not possible with a conventional atomic force microscope (AFM) due to the AFM’s Z-axis travel range limitations, and AFM restricted working space making it unable to accommodate large size of tissue and bioreactor. In addition to direct measurements, we also conducted finite element analysis simulations. Both experimental and numerical results point to the existence of such a correlation.

Figure 1.

Figure 1.

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Micromachined force sensor and human engineered cardiac tissue set up. a) Schematic diagram of the experimental setup. A Wheatstone bridge circuit with one of the resistors being the strain gauge (force) sensor of the microcantilever, the voltage signal is amplified and fed into a digital acquisition (DAQ) card. On the bottom diagram, a cantilever is shown in contact with the tissue. The vertical and axial force directions are schematically shown, b) Picture of human engineered cardiac tissue (hECT) on flexible end-posts.

2. MATERIAL AND METHODS

2.1. Micromachined force sensors and interface circuitry

A specialized micromachined cantilever force sensor was fabricated for these measurements similar to the ones published previously (Fig. 1a) to measure contractile forces [26, 27]. The sensor operates like an electronic finger, detecting contact with a cell and registering cellular or tissue contractions. We chose polyimide as a structural material because sensors made from polyimide offer high compliance, sensitivity, robustness, biocompatibility, and durability [26, 27]. The spring constant, which is the measure of the compliance of the cantilever, is given by: KN = E w t3 / (4 L3), where E is the Young’s modulus, L is the length, t is the thickness, and w is the width. Using polyimide with Young's modulus 3.2 GPa, in place of conventional MEMS materials such as Si3N4 80 GPa and Si 110 GPa, the device offers a low spring constant and as a result it is less damaging to tissue measurements. The cantilever sensor includes a thin film strain gauge made from Cr/Au with 1 nm/10 nm thickness that is sandwiched between two layers of polyimide, 1.5 μm and 0.5 μm, respectively. The sensor is 40 μm wide, 150 μm long, and approximately 2 μm thick. The fabrication process is implemented using micro-electromechanical systems (MEMS) methods and requires four masks. The detailed fabrication process flow is shown in Figure 2. In brief, a thermal oxide mask is deposited and patterned (1). The bottom layer of polyimide is spin-coated and patterned with reactive-ion etching (RIE) (2). A thin Cr/Au (1nm/10nm) metal line is sputtered for the resistor and a thicker Cr/Au (1 nm/100nm) is deposited on the pad area for electrical bonding (3). The top layer of polyimide is spin-coated and patterned with RIE (4). The cantilever structure is formed by a backside deep reactive-ion etching (DRIE) process (4). The oxide acts as an etch stop. Finally, the thermal oxide layer is removed in buffered oxide etch (BHF) (6). Cr/Au with 5 nm/25 nm is deposited on the backside of the wafer for refection. The resulting device is shown in Figure 2b. The inset of Figure 2b shows the device chip, which includes two sensor cantilevers one of which is broken off. The device chip is wired and passivated using polydimethylsiloxane (PDMS). The sensor’s spring constant, k, is ∼0.059 N/m and its nominal resistance is between 160 and 170 Ω.

Figure 2.

Figure 2.

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Micromachined force sensor fabrication, a) Schematic of fabrication steps to produce the micro-cantilever force sensor, b) Scanning electron micrograph of the cantilever device (Inset: chip with two devices, one of which is broken off).

The movement of the cantilever is controlled with a motorized micromanipulator with 25 × 25 × 25 mm3 range and 0.1 μm minimal increment step (KT-LS28-MV, Zaber Technologies). Lab VIEW software was used to control the movement and record the output of the sensor. The electrical signal from the cantilever was measured using a Wheatstone bridge circuit connected to a low noise amplifier, and a National Instruments digital acquisition (DAQ) card to record the data into a computer (Figure 1a). The cantilever was calibrated by bringing it in contact with a hard surface such as a glass slide, and recording the change in voltage from the bridge circuit output with distance travelled in the Z-axis. Then, the force can be estimated by multiplying the spring constant of the microcantilever with the distance travelled.

2.2. Human engineered cardiac tissue fabrication and optical tracking of post deflection

The induced pluripotent stem cell (iPSC) line (SKiPS-31.3) was derived from adult dermal fibroblasts, which were obtained with informed consent from a healthy adult [31]. This research project involved only secondary analysis of already collected biological samples without personal identifiers. iPSCs were grown in StemFlex media on tissue culture 6-well plates coated with hESC-qualified Matrigel in 5%CO2 at 37°C. iPSCs were differentiated into CMs by following a monolayer-based differentiation protocol [30]. Briefly, when iPSCs reached ∼80% confluence, StemFlex media was replaced with basal medium (RPMI 1640 media plus B27 supplement minus Insulin), and 10 μM CHIR99021 for 24 hours. On day 1, the media was replaced with basal medium without CHIR99021 for two days. On days 3 and 4, media was replaced with basal media with 5 μM IWR-1. On day 5, the media was replaced with basal media. On day 7, basal medial was replaced with RPMI 1640 media plus B27 supplement. Afterward the media was exchanged every 2 days. Functional iPSC-CMs appear in culture between days 7 and 10 post-CHIR treatment.

For hECT fabrication, iPSC-CMs were collected after 24 days of differentiation, by enzymatic treatment of the CM monolayer (using trypsin 0.25%); the CM cell pellet was resuspended in a solution of type-I collagen and Matrigel, 100ul of this cell-matrix mix was layered into the well of a single-tissue bioreactor [30] with each resulting hECT containing approximately 1.5 × 106 cells. The bioreactor was placed in a 60mm dish and after 2-hour incubation for polymerization, the bioreactor was covered with culture medium (RPMI 1640 media plus 2% B27 supplement), and maintained at 5% CO2 and 37°C with half-media exchanges every other day.

The single-tissue bioreactor is fabricated of polydimethylsiloxane (PDMS) with flexible end posts, the posts serve as anchors during formation of the hECT, and act as integrated force sensors that deflect during hECT beating [30, 32]. The average Young's modulus of the posts was 1.33 MPa from calibration tests using a high-sensitivity force transducer (Scientific Instruments, Heidelberg, Germany). The hECT develops in the shape of a tissue strip, held between the flexible end posts (Figure 1b). The hECT typically begins to beat spontaneously within 72 hours of fabrication, and deflection of the end posts with each contraction can be visualized with low magnification microscope by day 6. The axial force of the engineered tissues is measured using an established method to optically track end post deflections for force calculation described in [30]. The contractile function of each hECT was measured in a laminar flow hood in a custom setup that allows optical tracking of PDMS post deflection versus time using custom Lab VIEW software to acquire real-time data sampling (Figure 3). The data is then analyzed with a custom MATLAB script to calculate twitch parameters including developed force (DF; calculated from the post deflection with each twitch), and developed stress (DS; DF divided by hECT cross sectional area), as previously described [30]. Recordings were first performed without electrical stimulation to obtain data on spontaneous beating frequency. Then the hECTs were electrically paced by field stimulation (12-V biphasic pulse with 5-ms duration) at increasing frequencies, using a programmable Grass S88X stimulator (Astro-Med, West Warwick, RI).

Figure 3.

Figure 3.

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Optical tracking of post deflection for axial force calculation, a) Screen view captured during live tracking of post-displacement during hECT contraction. Image shows thresholded view of hECT, focusing on the top of each of the posts (black arrows); the inward post-displacement during each contraction (red arrows) is captured, and simultaneously, the difference in distance between the posts is displayed on the screen across time; with each contraction (red arrows) the distance between the posts decreases (white arrow), b) Side view of the hECT (with superimposed schematic on the left post) to show analysis of the post deflection based on the elastic beam theory (formula displayed in the inset), allowing calculation of the tissue force applied at some distance from the top of the post where deflection is measured. F is tissue contraction force; E, R, L represent Young's modulus, radius, and length of the posts, respectively; a is the height of the tissue on the post, which also corresponds to the point where it is estimated that the force is applied; δ is measured post deflection. Superimposed schematic on right post to show that the width of the contact area between the cardiac tissue and the post is measured (blue line) and the point where the force is applied is calculated halfway along the width of the tissue (red arrow).

2.3. Experimental set-up to use the microcantilever for vertical force measurement

The microcantilever was positioned in close proximity to the hECT and brought in contact with the tissue under optical observation using the motorized stage (Figure 4a, b). Due to the large range of motion afforded by the system, it is relatively simple to achieve contact with the sample. While moving the microcantilever closer to the sample the voltage signal coming out of the bridge circuit is monitored. A sudden increase in the voltage is indicative of contact. Once in contact, contractions start being recorded as shown in Figure 4c; the voltage changes of the microcantilever are monitored and plotted using Lab VIEW (a visual programming language). The vertical force measured by the microcantilever are compared to the axial force measurements to determine the relationship between these two forces.

Figure 4.

Figure 4.

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Vertical force measurement set up. a) Picture of microcantilever immersed in media and in contact with hECT. b) Picture of microcantilever approaching the hECT. c) Screenshot of real time LabView recording of microcantilever voltage changes as the tissue in contact contracts.

2.4. Measurements of elastic modulus

To measure the tissue’s elastic modulus, we removed the hECT from its bioreactor, and glued it down with Kwik-cast sealant (World Precision Instruments) to the bottom of a petri dish, applying the sealant to polar ends of the hECT, after 1 minute the dish was filled with media for subsequent measurements under AFM. The AFM measurements were performed with Flex-Bio AFM (Nanosurf AG, Switzerland) equipped with an inverted microscope Axio Observer (Carl Zeiss) and an environmental control enclosure. The AFM was operated in static force mode. The cantilever had a spring constant of 0.27N/m and silicon tip radius < 10nm (Stat0.2LAuD, Nanosurf AG, Switzerland). We measured force spectroscopy at 13 different spots (N=8 for each spot) with distances from the edge of the tissue ranging from 200 to 300 μm as shown in Figure 5a. The force curves, as shown in Figure 5b, were analyzed and the elastic modulus was obtained by applying the Hertz model [33, 34] to the forward approach curves using the ANA control and analysis software (Nanosurf AG, Switzerland) resulting in an average elastic modulus of approximately 10.7 +/− 3.7 kPa for the hECT. This value was later used in the finite element (FEM) model using COMSOL Multiphysics (ver. 5.4, COMSOL, Inc., MA, USA).

Figure 5.

Figure 5.

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Atomic force microscopy measurements, a) Optical image of AFM cantilever on tissue. The tip of the cantilever is about 250 μm from the edge of the tissue, which is within the mid segment of the width of the hECT. b) A sample force curve taken by AFM. The blue trace is the force curve on approach, and the orange trace is the force curve on retract. We observe hysteresis due to deformation and adhesion of the tissue during retract

2.5. Simulations

The tissue was simulated as a linear elastic material. The “Structural Mechanics” module in COMSOL Multiphysics was used to model the elastic behavior of hECT using an Isotropic Solid model (Figure 6). For the simulated tissue a Young’s modulus of 11 MPa, a Poison ratio of 0.49, and a density of 1081 kg/m3 were used. An initial force/unit area of 56.7 Pa was applied on each side of the tissue which is similar to the force measured using the post optical tracking method [30]. The force was reduced in 4 Pa steps in a parametric sweep. The posts material used was PDMS with density 970 kg/m3, Young's modulus 1.3 MPa, and Poisson's ratio 0.49.

Figure 6.

Figure 6.

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Simulation of hECT. Model of tissue on PDMS posts.

3. RESULTS AND DISCUSSION

We performed the experiments under electrical stimulation. The hECT beat spontaneously and contractile parameters can be recorded without electrical stimulation; however, the hECT display a force-frequency relationship [28, 35], therefore, the best means to do a comparison of force sensing systems is to do it at prescribed beating frequencies. Therefore, the tissue was paced between 0.2 and 1 Hz. The cantilever accurately detected the beating frequency of the hECT at different pacing frequencies (Figure 7a–e). The magnitude of the force measured by the microcantilever (vertical force) drops with increasing frequency, which is consistent with observations using optical tracking of the post deflection (axial force). The force-frequency relationship is linear at small frequency values below 1 Hz (Figure 7f and Supplemental Figure S1). The force displayed by the hECT is frequency dependent, this characteristic has been described by us and others [36–40]. Linear regression was used to examine the association between force and frequency, which is described by the equation: Force per unit area = −0.23 frequency + 0.27 (R2 = 0.99). The results demonstrate that the microcantilever can measure frequency and force in response to pacing hECT. At zero frequency, there is still a force, which points to the existence of initial stress. This initial stress is due to the contact of the microcantilever with the tissue during the approach step.

Figure 7.

Figure 7.

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Force plots of hECT obtained with micromachined force sensor at multiple frequencies, (a-e) In the Y-axis the cantilever reading in volts from 0–0. IV This is the reading from the Wheatstone bridge circuit that was amplified × 1000. In the X-axis the time is plotted in seconds, (f) Plot of force per unit area with frequency. At this frequency range the change is highly linear. As the frequency increases the amplitude of the contraction and the force decrease. (N= 6, error bars represent SD).

We then examined the association between the axial force measured from the “post deflection” force (Y-axis of Figure 8) and the vertical force measured with the cantilever (X-axis of Figure 8) using a linear regression model. A significant regression equation was found (axial force (1, 3) = 169.5, p < 0.001) with R2 of 0.98. The line of best fit is described by the equation: axial force = to 0.071 + 0.248 vertical force. Thus, the vertical force is a significant predictor of the axial force. Our experimental findings suggest that the axial and the vertical force are strongly correlated. These results are significant because they establish a relationship between axial and vertical force measured by the micromachined cantilever. These results also indicate the existence of initial stress in the axial force, which is observed in hECT since the posts are bent at a rest position (the tissue compacts overtime).

Figure 8.

Figure 8.

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Comparison of axial (post force) and vertical force (cantilever force). Plot of “post deflection” force in the y-axis vs. the cantilever force measured in the x-axis. Inset include linear regression model.

To develop a simulation model, the modulus of elasticity of the tissue needed to be measured. Using the AFM force curves measurement, described previously in Figure 5b, the modulus of elasticity was calculated to be 10.7 +/− 3.7 kPa. In the simulation, an approximate value of 11 kPa is used. Using the simulation model shown in Figure 9a the movement in the Z-axis of the tissue at various locations was calculated. The distance in the Z-axis changes in a parabolic manner when a force is applied on the sides of the tissue emulating the hECT movement. A distance of 1mm from the post’s outer edge was chosen to examine the vertical Z displacement due to the axial force applied in the X-axis. The Z-axis displacement is converted into force per unit area using the microcantilever spring constant and assuming a contact area of 30 μm × 40 μm. As shown in Figure 9b the relation between the vertical force (measured by the cantilever) and the axial force is linear and with the same order of magnitude as the physical measurements. Furthermore, the vertical displacement and therefore the force detected by the cantilever depends on the location the cantilever is placed on the tissue. The closer to the post the smaller the vertical force.

Figure 9.

Figure 9.

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Finite element analysis simulations, (a) Simulation of tissue bent due to force applied to both sides (using a scale factor of 10 to make the bending visually more pronounced), (b) Plot of “post deflection” force in the y-axis vs. the cantilever force measured in the x-axis at a distance 1 mm from the outer edge of the post. Inset includes the line of best fit and equation.

Finite element analysis (FEA) simulations provide a good approximation to the results obtained using the microcantilever. However, there are several differences between the experimental conditions and the FEA. First, the engineered tissue is not placed exactly on the top part of the post, in contrast the model assumes that the top part of the tissue and the post are on the same level. Second, there are tissue-to-tissue variations in elasticity, size, and shape. Third, the displacement of the cantilever and therefore the force measured is dependent of the location the cantilever is placed, as noted above. Finally, in the experimental measurement the cantilever has an initial stress as it is brought in contact with contracting tissue.

Prior publications have reported measuring single cell vertical force measurements using AFM techniques and other force transduction methods [24, 26, 41–43]. Indeed, when evaluating cardiomyocyte monolayers, we showed that the vertically measured beating frequency, captured by the movement of a similar micromachined cantilever, correlated with axial measured frequency and movement using optical tracking [26]. In other studies, our group and others have inferred single cell behavior from tissue measurements by dividing the axial force exerted by the hECT to the total number of myocytes used in the fabrication of the hECT [32, 44]. The structural composition of the hECT is characterized by the presence of CMs distributed throughout the hECT (Supplemental figure S2). The cells are organized as a syncytium within the hECT and this is reflected by the uniform contraction of the hECT with each electrical impulse. With each contraction of the hECT, the flexible posts bend inwards (Supplemental video 1). Although with limitations, we simulated this behavior with finite element analysis. While the microcantilever performs a local tissue measurement and the optical tracking reflects entire tissue behavior, it is highly encouraging that the contractile behavior that we detected using the optical tracking method correlates with the microcantilever measurements. Our observations demonstrate clearly this correlation when it was tested in hECT of cylindrical shape and that is suspended between flexible posts; but we cannot establish that this correlation applies in the study of single cells or other 3-D structures that display a contractile behavior. Further studies will be required using CMs to test this correlation of axial and vertical force at the single cell level.

4. CONCLUSIONS

Our results show that the microcantilever force sensor could detect reliably the frequency and vertical force of the hECT. The implementation of the microcantilever force sensor for measuring force of the hECT was used in this study with the purpose of investigating how vertical and axial force compare. Our gold standard for the measurement of force of the hECT is the optical force tracking system, which effectively provides a measurement of axial force, remains as an efficient and practical method for hECT contractile force analysis. However, at the single cell and monolayer level, iPSC-CMs force analysis would benefit from the implementation of a vertical force detection system. While preliminary, we have established, through experimental results and finite element analysis simulations, the evidence of correlation of axial and vertical force for the study of active contractile force. In the future, we plan to improve upon existing microcantilever systems for single cell vertical force measurement. This approach will facilitate acquisition of contractile properties of CMs, which are relevant for the evaluation of drug-induced cardiotoxicity effects.

Supplementary Material

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Download video file (513.7KB, m4v)

Highlights.

  • Novel micromachined sensor for measuring force contractions

  • Tool applicable to engineered tissue contractions and cardiomyocyte monolayer force

  • Correlation between axial force and vertical force established experimentally

  • Simulated behavior with finite element analysis comparing axial and vertical force

ACKNOWLEDGEMENTS

This work was supported by the National Institutes of Health (U24 DK112331-03S1, K01HL133424 and R03HL154286).

Abbreviations

AFM

atomic force microscopy

CM

cardiomyocyte

iPSCs

induced pluripotent stem cells

MEMS

micro-electromechanical systems

hECT

human engineered cardiac tissue

Footnotes

Declaration of interests

The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.

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