Analysis of Calcium Transients and Uniaxial Contraction Force in

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Analysis of calcium transients and uni-axial contraction force in single human embryonic stem cell derived-cardiomyocytes on microstructured elastic substrate with spatially controlled surface chemistries Eleonora Grespan, Sebastian Martewicz, Elena Serena, Vincent Le Houérou, Jürgen Rühe, and Nicola Elvassore Langmuir, Just Accepted Manuscript • DOI: 10.1021/acs.langmuir.6b03138 • Publication Date (Web): 19 Sep 2016 Downloaded from http://pubs.acs.org on September 29, 2016

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Analysis of calcium transients and uni-axial contraction force

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in single human embryonic stem cells derived-cardiomyocytes

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on microstructured elastic substrate with spatially controlled

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surface chemistries

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E. Grespan1, S. Martewicz2,3, E. Serena2,3, V. Le Houerou4, J. Rühe5, N.Elvassore2,3*

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1.

CNR Institute of Neuroscience, Corso Stati Uniti 4, 35127, Padova, Italy;

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2.

Department of Industrial Engineering (DII), University of Padova, Via Marzolo 9, 35131,

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Padova, Italy;

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3.

Venetian Institute of Molecular Medicine (VIMM), Via Orus 2, 35129, Padua, Italy

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4.

Institute Charles Sadron (ICS), University of Strasbourg, 23 rue du Loess, 84047, Strasbourg,

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France; 5.

Institute of Microsystems Engineering (IMTEK), Department of Chemistry and Physics of Interfaces, University of Freiburg, Georges-Köhler Allee 103, 79110, Freiburg, Germany

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* Corresponding author:

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Prof. Nicola Elvassore

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Department of Industrial Engineering (DII),

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University of Padova

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Via Marzolo 9, 35129

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Padova, Italy

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Tel: +39-049-827-5469

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Fax: +39-049-827-5461

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Email: [email protected]

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ACS Paragon Plus Environment

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Abstract

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The mechanical activity of cardiomyocytes is the result of a process called excitation contraction

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coupling (ECC). A membrane depolarization wave induces a transient cytosolic calcium

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concentration increase that triggers activation calcium-sensitive contractile proteins leading to cell

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contraction and force generation. An experimental setup capable of acquiring simultaneously all

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ECC features would have an enormous impact on cardiac drug development and disease study.

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In this work, we develop a micro-engineered elastomeric substrate with tailor-made surface

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chemistry to measure simultaneously the uni-axial contraction force and the calcium transients

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generated by single human cardiomyocytes in vitro. Micro-replication followed by photocuring is

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used to generate an array consisting of elastomeric micropillars. A second photochemical process

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is employed to spatially control the surface chemistry of the elastomeric pillar. As result, human

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embryonic stem cell derived cardiomyocytes (hESC-CMs) can be confined in rectangular cell-

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adhesive areas, which induce cell elongation and promote suspended cell anchoring between two

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adjacent micropillars. In this end-to-end conformation, confocal fluorescence microscopy allows

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simultaneous detection of calcium transients and micropillar deflection induced by single cell uni-

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axial contraction force. Computational finite elements model (FEM) and 3D reconstruction of cell-

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pillar interface allow force quantification. The platform is used to follow calcium dynamics and

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contraction force evolution in hESC-CMs cultures over the course of several weeks. Our results

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show how a biomaterial-based platform can be a versatile tool for in vitro assaying cardiac

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functional properties of single-cell human cardiomyocytes, with applications in both in vitro

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developmental studies and drug screening on cardiac cultures.

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Keywords

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Human embryonic stem cells derived cardiomyocytes

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Micropillar

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Photopattern

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Force generation

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Single cell

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1. Introduction

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Cardiac output is mainly determined by the functional performance of its elementary contractile

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unit: the single cardiomyocyte (CM) [1-3].

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At single-cell level, the events leading to force generation are comprehensively termed excitation-

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contraction coupling (ECC). This process involves voltage-gated ion channels opening, a rapid

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increase in intracellular calcium concentration and actin–myosin system activation by displacement

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of Ca2+-sensitive troponin complexes to generate cell shortening. Calcium acts as the second

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messenger linking excitation to contraction by entering the myocyte through voltage-activated L-

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type channels and triggering additional calcium release from intracellular stores via calcium-

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induced calcium-release (CICR) by ryanodine receptors (RyR) on the sarcoplasmic reticulum (SR).

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The initial conditions and mechanical relaxation are then restored by fast calcium clearance from

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the cytosol mainly by two routes: reuptake into the SR by sarcoendoplasmic reticulum Ca2+-

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ATPase (SERCA pumps) and extrusion through the sarcolemma by the Na+/Ca2+ exchanger

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(NCX) [4]. Thus, the centerpiece of the entire ECC is represented by the momentary calcium

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concentration change in the cytoplasm, called Ca2+ transient [5]. Deranged contractions of

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cardiomyocytes can cause dysfunctions in cardiac pumping activity and cardiac failure, which is

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currently the leading cause of mortality and morbidity in Western Countries [6]. For this reasons,

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detailed assessment of human cardiac function at the single cardiomyocyte level can help to

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elucidate the pathophysiological mechanisms of heart failure and facilitate development of novel

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therapeutic interventions.

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In recent years, human cardiomyocytes derived from pluripotent stem cells (hPSC-CMs) have

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arisen as one of the most promising cellular models for human myocardium for in vitro studies.

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Despite their immaturity and early phenotype, these cells are currently being employed in

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numerous studies as models for human cardiomyocyte function. Single hPSC-CMs have been

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widely assayed and characterized for their functional performance in electrical activity and calcium

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handling (7, 8). To assess both features, a plethora of methods have been employed for either

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high-content or high-throughput analysis, such as patch-clamp or micro-arrayed electrodes (MEA)

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for action potential (10, 11) and imaging techniques for calcium cycling (12, 13, 14). Nevertheless,

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fully comprehensive functional assays should not omit the evaluation of the end-point cardiac

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features along with the other elements of the ECC: force generation and contraction

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characteristics.

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Despite several experimental methods have been employed to quantify cellular forces, such as

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AFM [14-16], flexible cantilever [17], carbon fiber-based approaches [18], flat elastic substrates

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(traction force microscopy) [19-23], strain-gauges [24] and arrays of microposts [25-34], none of

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these provides a truly high-throughput method. Moreover, they are bound by experimental

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constrains resulting in difficult comparisons with in vivo conditions especially because in most

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cases isolated CMs are analyzed in a flat environment. In 2D cultures, the recorded tension forces

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are not generated by end-to-end cell configurations that characterize the functional syncytium of

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the cardiac muscle, but rather tensions are transduced from a nearly flat cell to the specific

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substrate of choice through adhesion of the whole cell membrane.

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Most reported studies are based on primary animal-derived cell models. In order to overcome the

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intrinsic differences between animal and human physiologies, a human model would be more

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desirable. In this perspective, the generation of human pluripotent stem cell-derived

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cardiomyocytes (hPSC-CMs) provides an excellent tool for studying in vitro cardiomyocytes of

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human origin in physiological conditions as well as in genetically or environmentally pathological

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ones.

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In force-measurement studies specifically involving hPSC-CMs, mainly two approaches have been

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used: traction force microscopy [35-37] and dense micropost arrays [38, 39]. In both cases,

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cardiomyocytes are cultured on top of a compliant substrate and the total force produced by a

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single cell is determined by summing the absolute magnitudes of the forces measured at each

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point of interest, evaluated by physical displacement of microbeads embedded in an hydrogel or

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the edge-movement of elastic microposts. These values represent tangential traction force of the

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cardiomyocyte on a widespread flat surface, thus the resulting sums can be scarcely approximated

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to two dipoles and are not equivalent to the gold standard measurement of axial traction [40].

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In order to overcome the 2D limitations, engineered human cardiac tissues hanging between two

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micropillars have been proposed as solution [41-43]. In these studies, the uni-axial traction force is

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measured in dense cellular agglomerates without single-cell resolution.

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A single-cell uni-axial approach has been recently proposed by using a thermoresponsive

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sacrificial support layer in conjunction with an array of widely separated elastomeric microposts

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[40]. Through this method, it was possible to achieve culture of immature cardiomyocytes,

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including human embryonic stem cell derived (hESC) cardiomyocytes, which randomly anchor

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between two or multiple pillars. However, the force was quantified only for neonatal rat

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cardiomyocytes by phase contrast imaging, without taking into account the relevant calcium

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dynamics data associated with each contraction and critical for the cardiac functionality

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assessment [5].

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In this work, we aimed at developing an in vitro platform for simultaneous acquisition of calcium

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transients and uni-axial contraction force at a single human cardiomyocyte level. In particular, we

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developed a photo-patterning technique on elastic micropillared substrate in order to guide the

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cells into a physiological end-to-end cell configuration for long-term functional evaluation of

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hESCs-CMs performances. Previous studies on cells in contact with micropillared surfaces [38-43]

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were based on polydimethylsiloxane (PDMS). A limit of PDMS substrates is its challenging surface

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chemistry, which hardly allows permanent chemical modification. To adapt to the specific biological

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aim, our platform required a novel polymer, which characteristics allow easy microstructuring and

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stable chemical modification, in order to finely tune both topological and chemical properties of the

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culture substrate. For substrate generation, we used a biocompatible elastomeric substrate

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consisting of n-butylacrylate (n-BA) and methacryloxybenzophenone (MABP) as a photo-active

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group. Microstructured surfaces composed of micropillar arrays were fabricated by replica molding

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and photocuring. This was followed by spatially controlled photochemical polymerization of linear

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polyacrylamide to define cell-adhesive and cell-repellent areas on the microstructured substrates

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with micrometric resolution. Fluorescence confocal microscopy together with computational

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mechanical modeling of three dimensional cell-pillar interface reconstructions provided

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quantification of calcium transient parameters (Ca2+ Rel and Ca2+ Up) and uni-axial contraction

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force (Figure 1).

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Figure 1: Functional characterization of single hESCs-CM. hESCs-CMs are seeded on elastomeric

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microstructured substrates that have been previously photopatterned to obtain cell-adhesive and

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cell-repellent areas on the surface.

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contraction. Confocal analysis allows the acquisition of calcium transients, recording of micropillar

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deflection and visualization of focal adhesion points, and are repeated after 1 week and after 5

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weeks of culture. The use of a proper finite element model (FEM) allows the quantification of

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contraction forces.

Cells anchor to the micropillars and bend them upon

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2. Materials and methods

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2.1 Substrate fabrication

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The elastomeric flat and microstructured substrates have been fabricated by a cell-adhesive

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photopatternable material consisting of n-butylacrylate (n-BA) and methacryloxybenzophenone

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(MABP). In particular, P(nBA-4%MABP) was obtained through free radical polymerization as

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described in the Supplementary Information and in previous publications [44-46]. To tune cell

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growth only in well defined regions, a thin layer of cell-repellent polyacrylamide (PAA) has been

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photoattached by UV-light exposure to the P(nBA-co-4%MABP) surface, which prevented protein

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adsorption and cell adhesion everywhere, except than in certain areas which were kept cell

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adhesive.

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2.1.1 Fabrication of PDMS stamps

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As first step, microstructured PDMS stamps have been realized by replica molding from micro-

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structured Si-wafers obtained through standard photolitography technique. Briefly, a Si-wafer was

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spin coated with a thin layer of negative photoresist (AZ-1518, MicroChemicals, Germany). The

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photoresist was illuminated by UV-light (λ=365 nm) through a chrome mask (Delta Mask, The

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Netherlands) that presented the desired pattern. Uncross-linked resist was removed rinsing the

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wafer in the proper AZ-1518 developer (MicroChemicals, Germany) for few seconds. The wafer

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was then etched to obtain the final micropillared stamp. The microstructured Si-wafer was

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subsequently silanized by exposure to the vapor of (tridecafluoro-1,1,2,2-tetrahydrooctyl)-1-

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trichlorosilane in vacuum for 30 minutes, to facilitate PDMS removal during the replica molding

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step. PDMS (Sylgard 184, Dow-Corning. Midland, MI) elastomer was thoroughly mixed with the

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silicone elastomer curing agent in a 10:1 ratio, poured over the microstructured Si-master and kept

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under vacuum for 1 h to allow the complete filling of the pattern and air bubbles removal. The

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sample was then cured at 80°C for 3 hours and subsequently peeled off the Si-wafers. These

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samples were used as stamps for the subsequent fabrication of P(nBA-co-4%MABP)

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microstructured substrates.

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2.1.2 Fabrication of P(nBA-co-4%MABP) substrates

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The fabrication of microstructured P(nBA-co-4%MABP) samples is described in figure 2: P(nBA-

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co-4%MABP) polymer was dissolved in toluene (300 mg/ml) and a 60 µl drop of the resulting

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solution was poured on the appropriate PDMS stamp. A silanized glass slide was then gently

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pressed on the top of this drop, allowing the still not cross-linked polymer to completely fill the

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pattern of the stamp. A covalent bond between the glass slide and the polymer was achieved

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thanks to the use of triethoxybenzophenone (3EBP) silane, as described in the Supplementary

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Information. This sandwich formed by PDMS stamp, P(nBA-co-4%MABP) and silanized glass was

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exposed to UV-light (λ = 365 nm, P = 6.5 mW ) for 20 minutes. The cross-linked microstructured

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samples covalently bonded to the glass slide were then peeled-off from the PDMS stamp.

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To obtain flat P(nBA-co-4%MABP) samples, the solution was spin-coated on a silanized glass at a

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velocity of 600 rpm for 1 minute and then irradiated by UV-light.

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2.1.3 Photopatterning of flat and microstructured P(nBA-4%MABP) substrates with PAA

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Acrylamide solution was prepared diluting acrylamide (Sigma Aldrich, Germany) in deionized water

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obtaining a final concentration of 8%. PAA was supplemented with 30 mg/ml of photoinitiator

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(Irgacure, BASF, Germany) previously diluted in 100 µl methanol. Figure 2B shows the steps

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necessaries to photopattern the P(nBA-4%MABP) substrates: A 70 µl drop of acrylamide/Irgacure

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solution was poured on a glass coverslip, the P(nBA-co-%MABP) sample was turned upside down

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and brought into contact with the polyacrylamide solution, and a photomask was gently placed on

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the glass that supported the P(nBA-co-%MABP) sample.

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This sandwich-like sample was then irradiated for 90 s with UV light (λ = 365 nm). As a result of

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this process in the irradiated areas a surface-attached polymer network was generated. After UV

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light exposure the whole elastomeric sample was covered by a thin layer of cell-repellent PAA,

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except those areas that have not been shaded by the photomask and thus have not been exposed

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to UV light. The sample was rinsed with deionized water for 3 minutes to remove all the residual

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not cross-linked PAA. The same procedure has been applied both on flat elastomeric samples and

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on microstructured elastomeric samples.

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2.2 Mechanical characterization of the substrate

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2.2.1 Dynamic Mechanical Thermal Analysis

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The elastic properties of P(nBA-co-4%MABP) were tested through Dynamic Mechanical Thermal

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Analysis (DMTA, developed on Instron 4502 Tensile Machine). To perform the DMTA test that

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allowed to quantify the Young’s modulus of P(nBA-co-4%MABP) at different temperatures, specific

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samples have been fabricated: the samples needed to be at least 3 mm thick and their length had

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to be at least twice bigger than their width. For these reasons samples were fabricated by casting.

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A solution of P(nBA-co-4%MABP) in toluene was prepared (300 mg/ml) and 2 ml of this solution

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were poured in a rectangular Teflon stamp (2x5 cm). The solution was then exposed to UV-light for

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30 minutes. The sample was peeled-off, turned upside-down, and exposed to UV-light for further

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30 minutes until the film was totally cross-linked. The film thickness ranged from 3 to 3.5 mm. A

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preload of 0.002 MPa was applied to each rectangular sample to assure the same tensile condition

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for all the different samples before starting the test and a tensile loading all along the test. A

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sinusoidal stress with a frequency of 1 Hz was applied deforming the samples of 1% maximum.

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The frequency was fixed at 1 Hz because this frequency is comparable to the frequency of

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deformation induced by the cells during contraction. The experiment was repeated for 5 samples.

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2.1.2 Nanoindention tests

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A nanoindenter (Nanoindenter CSM instruments, Switzerland) was used to experimentally

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reproduce the process of bending of the micropillars when a tangential force is applied to the pillar

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head. To this, a glass sphere was brought into contact with the microstructures of the specimen

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and dragged laterally to induce bending of the micropillars while the tangential force was recorded.

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When a specific tangential force is applied, the displacement of the top of the micropillars depends

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on the material properties (Young’s modulus) and on the geometry of the micropillar array. An in-

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situ camera that recorded the contact through the transparent specimen during the test allowed to

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simultaneously acquire the tangential force applied by the instrument and the micropillar deflection

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(Figure 3B). More details on the nanoindenter test are provided in the Supplementary Information.

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2.3 Finite element modeling

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Finite element model (FEM) of the substrate was realized using the MSC MARC software. The FE

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model of the substrate was developed according to the effective geometry configuration. Because

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of the simple tensile response of P(nBA-co-4%MABP) in the range of moderate strains, a Hookean

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constitutive model was adopted to model the material. Since the investigated cardiomyocytes act

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between single couples of pillars, it was possible to estimate contractile forces of cells through an

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indirect method, by applying assigned values of deflection (deduced experimentally) to the pillars

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and estimating the corresponding forces. Deflections of the pillar were applied as imposed

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displacements in the FEM. The force can be applied at different heights, as it happened in the

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case of cell measurements.

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2.4 Cell culture

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The cardiac differentiation protocol has been adapted from Lian et al., [50]. The hESCs were

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initially cultured on 0.5% Matrigel-coated (Corning) plates in mTeSR1 medium (StemCell

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Technologies) until confluence. Differentiation was induced by medium change to RPMI 1640/B-27

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w/o insulin (Invitrogen) supplemented with 12 µM CHIR99021 (Miltenyi). After 24 h, the medium

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was changed to RPMI 1640/B27 w/o insulin for another 2 days. The medium was then changed to

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RPMI 1640/B27 supplement w/o insulin and 10 µM IWP-4 (Stemgent). At day 5, the culture

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medium was switched again to RPMI 1640/B27 w/o insulin for 2 days and then the cultures were

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maintained for the entire culture time in RPMI 1640/B27. Medium was changed every 4 days. For

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disaggregation, a digestion mix solution was prepared with collagenase I (2 mg/ml), collagenase IV

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(1 mg/ml), DNase I (2 µl/ml) and 10 µM Y-27632 in PBS with Ca2+ and Mg2+. Cells were incubated

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for 25 minutes at 37°C in this solution, washed in PBS w/o Ca2+ and Mg2+ and digested to single-

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cell suspension with TryPLE Select reagent (Life technologies).

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Single cell suspensions were plated on the substrates coated for 1 hour with 20 µg/ml laminin (BD,

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Bioscience, USA) at a 4000 cells/mm2 density. Cells begun to contract after 3-4 days of seeding

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and could be kept viable up to 8 weeks, changing the medium every 4-5 days.

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2.5 Immunohistochemistry

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Cells were fixed with PBS containing 2% paraformaldehyde (Sigma-Aldrich) for 7 minutes,

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permeabilized with PBS containing 0.5% Triton X-100 (Sigma-Aldrich), and blocked in PBS

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containing 2% horse serum for 45 minutes, at room temperature. Primary antibodies were applied

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for 1 hour at 37 °C. Cells were washed in PBS (Life Technologies) and incubated with

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fluorescence-conjugated secondary antibodies against mouse for 45 minutes at 37 °C. Finally,

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nuclei were counterstained with DAPI (Sigma Aldrich), and samples were viewed under Leica TCS

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SP5 fluorescence confocal microscope (Leica Microsystems, Wetzlar, Germany). Primary

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antibodies used were the following: mouse monoclonal anti-cTnT (Sigma Aldrich; 1:100 dilution),

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mouse monoclonal anti-vinculin (Sigma-Aldrich; 1:100 dilution). Secondary antibody used was goat

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anti-mouse (Invitrogen; 1:200 dilution). All antibodies were diluted in 3% bovine serum albumin

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(Sigma Aldrich).

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2.6 Calcium transients and micropillar deflection acquisition through confocal analysis

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hESCs-CMs were loaded in serum-free Dulbecco’s modified Eagle medium (Invitrogen)

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supplemented with 2.5 µmol/l of fluorescent calcium dye Fluo-4 AM (Invitrogen) for 20 minutes at

13

37°C in the presence of 2 µmol/l of Pluronic F-127 (Life Technologies) and 20 µmol/l of

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sulfinpyrazone (Sigma-Aldrich), then incubated for additional 10 minutes at 37°C without Fluo-4

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AM, and added with 0.2 µmol/l of di-8- ANEPPS (Invitrogen). Cell dynamics were obtained in

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recording solution: NaCl, 125 mmol/l; KCl, 5 mmol/l; Na3PO4, 1 mmol/l; MgSO4, 1 mmol/l; CaCl2,

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2 mmol/l; and glucose, 5.5 mmol/l, to pH 7.4 with NaOH. Line scans were acquired with a Leica

18

TCS SP5 fluorescence confocal microscope using a 63x oil immersion objective, with 488 nm Ar

19

laser line as an excitation source and 400 Hz acquisition frequency. Line scans were then

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analyzed using Matlab (MathWorks) software to obtain calcium transient profile and quantify

21

micropillar deflection. For the calcium release phase (Ca2+ Rel), the time to peak value was

22

calculated considering the time from baseline to the 95% of the maximum value of fluo-4 intensity

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recorded by the confocal microscope. For evaluating the calcium reuptake rate after contraction

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(Ca2+ Up), the half-life of the calcium decay was considered. More information on the calculation of

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Ca2+ Rel and Ca2+ Up are provided in the Supplementary Information (figure S1). The total time of

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the calcium transients (calcium transient duration) has been calculated adding the time of calcium

27

release phase to two times the calcium reuptake phase: Ca2+ Tot = Ca2+ Rel + 2Ca2+ Up. The

28

deflection of the micropillars was recorded by following the fluorescence of di-8-ANEPPS, which

29

labeled both cell membranes and the surface of the micropillars, thanks to its lipophilic properties.

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2.7 Statistical analysis

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Data are presented as means +/- standard deviation. Data pairs were compared by non-directional

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Student’s t-test. Correlation coefficients were calculated through the Pearson’s method. All data

34

manipulation and computation was performed with the Matlab (MathWorks) software.

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3. Results

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3.1. Sample preparation

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The micropillar structures (size 10µm x10µm, height 20µm, distance 25µm) are generated through

3

a standard micro-replication process. However, instead of using a standard polymer we employed

4

a polymer containing benzophenone units incorporated into the polymer. Upon UV-irradiation

5

these groups become activated into a biradicaloid triplet state, which lead to linking of two polymer

6

chains through (formal) C-H insertion reactions [44-47]. To locally modify the surfaces a photopoly-

7

merization reaction of acrylic acid (AA) was carried out during brief UV-exposure, where a

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photomask prevented photoreaction to occur in the shaded areas. Transfer to polymer and

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recombination lead a crosslinking of the forming polymer chains and thus to the formation of a PAA

10

network. The forming PAA network become attached to the surface presumably in two ways: on

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the one hand grafting could occur through benzophenone units located at the surface, which had

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been “surviving” the photocuring of the pillar material, could become activated and bind to the PAA

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network. On the other hand, macroradicals could attach to the polymer substrate in a conventional

14

transfer reaction, where the macroradicals reacted with the substrate polymer. The thickness of the

15

PAA layer was determined by atomic force microscopy (AFM) and measured to be 250 ± 43 nm

16

(data not shown). The unmodified P(n-BA-4%-MABP) network surface is strongly non-polar and

17

thus allowed the adsorption of proteins such a laminin from solution rendering it cell adhesive. The

18

PAA-coated areas, however, represent a neutral, strongly swollen hydrogel, which prevented

19

protein adsorption and consequently cellular adhesion through entropic shielding/size exclusion

20

[47]. Thus cells brought into contact with such surface were strongly confined to that areas, which

21

were not covered by the hydrogel coating.

22 23 24

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Figure 2: Substrate fabrication. A) Fabrication of elastomeric microstructured substrates of P(nBA-

3

co-4%MABP): the polymer is crosslinked through UV light exposure. Upon illumination

4

benzophenone forms a biradicaloid triplet state, which abstracts a hydrogen atom from a

5

neighbouring C-H group of a polymer. The obtained closely adjacent two carbon radicals can

6

recombine and establish a covalent bond between the polymer chains and eventually crosslinking.

7

B) Photopatterning of PAA on P(nBA-co-4%MABP): PAA is exposed to UV light through a

8

photomask and photoattaches to P(nBA-co-4%MABP) surface forming a surface-attached network

9

only on the areas exposed to light. C) Laminin adsorption: laminin solution is incubated for 1 hour

10

on the substrate and the protein adsorbs on the P(nBA-co-4%MABP) surface.

11 12

3.2 Mechanical characterization

13

In order to mathematically derive the contraction force generated by the cells anchored on

14

micropillars, the determination of the mechanical properties of the elastomer employed in terms of

15

the modulus of the material is essential. Young’s modulus was measured by DMTA across an

16

interval of temperatures ranging from -40°C to 40°C (Figure 3).

17

applications such as cell cultures, temperatures of interest are above 25°C and the elastomer

In particular, for biological

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P(nBA-4%MABP) displayed a Young’s modulus of 1.5 ± 0.3 MPa at 36.5°C, corresponding to the

2

cell culture maintenance temperature (Figure 3C).

3

The bending tests performed on microstructured P(nBA-co-4%MABP) samples were instrumental

4

for developing a finite element model through which shear force applied to the micropillars and the

5

pillar deflection can be connected with each other. The samples consisted of a lattice of square

6

micropillars having a height o 10 µm, a width of 10 µm and an interaxial distance of 20 µm. A

7

round glass tip with a radius of 500 µm was put in contact to the top of the microstructured

8

substrate, without applying a normal load to avoid compression of the micropillars. Then the tip

9

was slightly moved laterally, and the tangential force produced between the tip and the bended top

10

of the elastomeric micropillars was recorded by the instrument while the camera showed the

11

displacement of the top of the micropillars. Assuming an equally distributed force among all the

12

micropillars in contact since the radius of the glass tip is much bigger than the micropillar width, the

13

tangential force ft acting on a single pillar could be calculated from the tangential force recorded by

14

the instrument (Ft) divided by the number of micropillars in contact. For the determination of the

15

deflection, the algorithm employed calculates the position of the centroid of the top of each pillar in

16

contact both in the undeformed (cundeformed ) and deformed configuration (cdeformed ). For each pillar

17

the displacement ∆xi is calculated as ∆xi = cdeformed − cundeformed. The final value of the displacement

18

d is given by the mean value of the displacements calculated for each micropillar: d = ∑ ∆xi / 9 for

19

i=1,..,9 .

20

The FE model has been validated by predicting the tangential forces which are must applied to the

21

top of the micropillars in order to achieve a certain deflection. Experimental data show a linear

22

elastic behavior of the pillar material for displacements of the top of the micropillars (∆x) of up to

23

2.5 µm. Within this range of displacement good agreement between experimental data and FEM

24

can be observed (Figure 3C). No permanent deformation was recorded once the force was

25

removed. Based on literature data [33,37], the expected contraction forces exerted by

26

cardiomyocytes on the micropillars fall within the linear response range of the plot, thus a Hookean

27

model for micropillar bending could be assumed.

28

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Figure 3: Mechanical characterization of P(nBA-co-4%MABP). A) Scheme of the bending test on

3

microstructured P(nBA-co-4%MABP) substrates. The glass sphere of a nanoindenter applies a

4

tangential force on the top of the elastomeric micropillars, bending them laterally (A.1) while

5

micropillar deflection is visualized through an in situ camera (A.2). Scale bar is 20 µm. B)

6

Tangential force vs micropillar deflection according to the experimental data obtained from the

7

bending test (circles) and the finite element model (line). C) Results of the DMTA test on P(nBA-

8

co-4%MABP) performed at a frequency of 1Hz. The Young’s modulus at 36.5°C was found to be

9

1.5 ± 0.3 Mpa.

10 11

3.3 Microscopy of the microstructures

12

The substrates presented in this work have been fabricated with the cell-adhesive elastomeric

13

polymer P(nBA-co-4%MABP), with a cell-repellent layer of polyacrylamide (PAA) photoattached to

14

the P(nBA-co-4%MABP) surface by photopolymerization with UV light.

15

patterns of PAA have been achieved by interposition between the sample and the light source a

16

chrome photomask, in order to cover and protect from PAA polymerization area desired to be cell-

17

adhesive (Figure 4A).

18

The cell-adhesive areas have been designed as rectangles with dimensions of 30x15 µm providing

19

a 2:1 ratio for cell adhesion and spreading. Moreover, as the micropillars have a width of 10 µm, a

20

height of 20 µm and an interaxial distance between pillars of 25 µm, these dimensions allowed to

21

coat such that the surfaces of two micropillars and the area between them remained cell adhesive,

Specific geometrical

13


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while the surrounding become strongly cell repellent (Figure 4B). The presence of the surface-

2

attached hydrogel in Figure 4B can be nicely seen through some wrinkling of the hydrogel occuring

3

during drying of the sample, which produces a clearly visible surface pattern.

4 5

6 7

Figure 4: Optical micrographs of surface-modified P(nBA-co-%MABP) substrates. Bright field

8

images of (A) a flat substrate with locally generated PAA layers and (B) of the same structures as

9

in (A) but on a microstructured substrate. Cell adhesive areas are marked by white dashed lines.

10

Scale bar is 30 µm. Details of the layer generation are given in the text.

11 12

3.4 Cell and substrate integration

13

hESCs-CMs were seeded on flat and on microstructured substrates after photopatterning with

14

linear PAA and functionalization with laminin solution. As the protein cannot adsorb to the PAA

15

coated areas, selective hESCs-CMs adhesion to the non-illuminated areas could be observed on

16

both types of substrates (Figure 5). hESCs-CMS could adhere to the cell-attractive areas of the

17

surface and after 4 days of culture the cardiomyocytes displayed spontaneous contracting activity,

18

demonstrating that laminin coated P(nBA-co-4%MABP)-surface does not impair cell function. The

19

cardiac culture could be maintained viable and contracting up to 8 weeks.

20

To evaluate the sarcomeric organization of hESCs-CMs, cells were fixed at day 20 to assay for

21

cardiac troponin-T (cTnT) by immunofluorescence staining (Figure 5B). Cardiomyocytes

22

constrained within the rectangular cell-adhesive areas spread assuming a rectangular shape and

23

displaying myofibril organization along the long axis of the cell (Figure 5A). When seeded on

24

microstructured and photopatterned substrates, hESCs-CMs anchored between two micropillars,

25

bending them during contraction (Figure 5B.1 and Figure 5B.2). The sarcomeric organization

26

perpendicular to the micropillars was assessed by immunofluorescence (Figure 5B.3), providing a

14


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clear proof of the generated force direction. Moreover, imaging of focal adhesion points by vinculin

2

staining (Figure 5B.4) allowed to determine at which height of the micropillars the force was

3

applied, in order to produce more accurate FEM data.

4 5

Figure 5: Cell and substrate integration. A) hESCs-CMs seeded on photopatterned substrates.

6

Cells selectively adhere to cell-adhesive areas (A.1) and assume a rectangular shape (A.2)

7

showing a sarcomeric organization comparable to that of embryonic cardiomyocytes, as indicated

8

by immunofluorescence analysis of cTnT (A.3). B) hESCs-CMs seeded on microstructured and

9

photopatterned substrates. Cells selectively adhere to cell adhesive areas extending between two

10

elastomeric micropillars (B.1) and deflecting them of the value ∆x upon contraction (B.2). B.3)

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Immunostaining of cTnT of a single hESCs-CM anchored between two micropillars. B.4) Focal

2

adhesion points indicated by vinculin immunostaining on a single hESCs-CM anchored between

3

two micropillars.

4 5

3.5 Contraction force and calcium transients along cell maturation

6

Fluorescence confocal microscopy allowed live imaging of fast dynamics of simultaneous events

7

by taking advantage of the spectral properties of the employed fluorescent dyes. The

8

cardiomyocytes were loaded with fluo-4 green fluorescent dye used to assay calcium dynamics,

9

while the red lipophilic dye di-8-ANEPPS was employed for micropillar visualization (Figure 6).

10

Line-scan mode was employed to achieve high recording speed (frequency of acquisition: 400 Hz)

11

and providing experimental data of calcium transient dynamics through fluo-4 fluorescence

12

intensity and micropillar edge deflection acquired against time (Figure 6A). Prior to data analyses,

13

we evaluated the effect of fluo-4 loading on calcium dynamics parameters. Comparison of Ca2+ Rel

14

and Ca2+ Up values for different loading conditions resulted in no significant differences when the

15

dye concentration was varied keeping the incubation time constant (Figure 2SA). When calcium

16

transient parameters were evaluated keeping the dye concentration constant and varying the

17

incubation time (Figure 2SB), we found a statistically significant deviation only at longer dye

18

incubation times than the 20 minutes used here.

19

Taking advantage of the grid of the microstructured substrate, the position of each individual

20

analyzed cardiomyocyte could be pinpointed for subsequent immunofluorescence analyses, aimed

21

at determining the exact height of cell anchorage on the pillars. This parameter, acquired through

22

anti-vinculin immunofluorescence staining and confocal sectioning, was essential for the correct

23

computation of the mathematical model employed to derive contraction force values (Figure 6B).

24

16


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Figure 6: Detection of calcium transient and quantification of contraction force on single hESCs-

3

CMs. A) Live-cell confocal analysis. A.1) Confocal image of a single hESCs-CM anchored between

4

two micropillars. Fluo-4 (green signal) allows to detect calcium transients while di-8-ANEPPS

5

adsorbed to micropillar walls (red) allows to detect micropillar deflection. A.2) Image of the

6

simultaneous acquisition of calcium transient (green signal) and micropillar deflection (red lines).

7

A.3) Plot of calcium transient (green) and consequent micropillar deflection (red). B) Quantification

8

of the contraction force from micropillar deflection. B.1) Confocal 3D image of a hESCs-CM

9

anchored between two elastomeric micropillars. After immunostaining of the cell with vinculin it is

10

possible to determine the location of the focal adhesion points of the cell, corresponding to the

11

points of application of the force in the model. B.2) FE modeling of the distribution of the tangential

12

force exerted by the cell when the contraction force is applied at a height of 16 µm.

13 14

Extension of time in culture of hPSC-CMs has been frequently linked to improvement in cardiac

15

functional features in single cardiomyocytes [48, 49]. To assess if our platform could detect such

16

trend, we analyzed the calcium transients and contraction force of hESCs-CMs after 1 week and 5

17

weeks from seeding on the microstructured substrates (Figure 7 A). Within this interval of time, the

18

calcium transient significantly shortened, both for calcium release rate and calcium reuptake rate

17


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(Figure 7B). Calcium release time (Ca2+ Rel) decreased from 0.35 ± 0.03 s at week 1 to 0.19 ± 0.02

2

s at week 5, whereas the half-life of calcium decay (Ca2+ Up) decreased from 0.36 ± 0.03 s at week

3

1 to 0.16 ± 0.02 s at week 5. We observed similar trends in calcium transients recorded in hESCs-

4

CMs cultured on flat P(nBA-co-4%MABP) substrates with Ca2+ Rel decreasing from 0.38 ± 0.05 s

5

at week 1 to 0.31 ± 0.03 s at week 5, and Ca2+ Up from 0.50 ± 0.09 s at week 1 to 0.38 ± 0.05 s at

6

week 5. Cells cultured on microstructured substrates showed better functional performance with

7

shorter calcium transients at both 1 week and 5 weeks of culture. The generated force associated

8

with the contraction in cardiomyocytes cultured on microstructured substrates has been quantified

9

for the same cells in culture for 1 and 5 weeks. We observed an increase in the developed force

10

from 502 ± 122 nN at week 1 to 719± 101 nN at week 5 of culture (Figure 7C). Interestingly, we

11

found a negative correlation (R=-0.784) between the length of the calcium transient and the

12

developed contraction force with cardiomyocytes displaying shorter calcium transients exerting

13

higher force on the micropillars (Figure 7D). Nevertheless, analyzing our data points for correlation

14

between Ca2+ transient length (R=0.081 and R=0.104, respectively week 1 and week 5) or

15

contraction force (R=0.032 and R=0.091, respectively week 1 and week 5) against beating

16

frequency, no statistical correlation was found. Assuming the immature phenotype of hPSC-CMs

17

[8], this observation is consistent with reported literature data, making consideration over inotropic

18

and lusitropic effects difficult (Figure 7E).

18


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Figure 7: Analysis of single hESCs-CM calcium transients and contraction force after 1 week and

3

after 5 weeks of culture. A) Examples of simultaneous acquisition of calcium transients and

4

contraction forces of single hESCs-CM after 1 week (A.1) and after 5 weeks (A.2). B) Comparison

5

of calcium release (B.1) and calcium reuptake (B.2) of single hESCs-CMs analyzed after 1 week

6

and after 5 weeks of culture on flat photopatterned substrates (light gray) and on microstructured

7

and photopatterned substrates (black). C) Contraction force of single hESCs-CMs quantified on

8

cells after 1 week (light gray) and after 5 weeks (black) of culture on microstructured and

9

photopatterned substrates. Cells analyzed after 5 weeks from seeding developed stronger forces

10

and shorter calcium transients than cells analyzed after 1 week from seeding. Student t-test p

11

values, p ≤ 0.01 (*), p ≤ 0.001 (**), n=15. D) Maximum contraction force vs calcium transients of

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cells analyzed after 1 week and after 5 weeks of culture. E) Distribution plot of calcium transients

2

duration (E.1) and maximum contraction force (E.2) against beating frequency.

3

4. Discussion

4

In this work, we presented an in vitro platform based on a micro-engineered elastomeric

5

biomaterial suitable for assaying functional features of human embryonic stem cell-derived

6

cardiomyocytes (hESC-CMs) through optical imaging. Our main goal was to provide a viable

7

technology for contraction-force measurement in a human single cardiomyocyte, with simultaneous

8

acquisition of calcium dynamics. To this end, we developed a microengineered substrate

9

fabricated with a novel material, tuning its surface chemistry in order to achieve geometrically

10

constrained human cardiac cultures. In particular, we showed that calcium transients and

11

contraction force of single hESCs-CMs can be simultaneously acquired in cell cultures that can

12

span several weeks (up to 8 weeks). This time frame is suitable for both testing the long-term

13

effects of pharmacological treatment and analyze the progression of cardiac features upon

14

extended culture time.

15

Through proper biomaterial design, micro-fabrication and micro-patterning techniques, we

16

successfully achieved fine control of cell shape in an end-to-end cell configuration between two

17

micropillars. Previous studies are based on PDMS micropillared surfaces [38-43]. However PDMS

18

surface chemistry is difficult to modify permanently; atmospheric air plasma and argon plasma

19

oxidation can be used to alter the surface chemistry, adding silanol (SiOH) groups to the surface.

20

Nevertheless oxidized surfaces are stable for only short periods of times, independently from the

21

surrounding medium, i.e. regardless of whether the surfaces are stored in vacuum, air, or water.

22

To overcome PDMS limits and obtain a surface having tunable surface mechanical and chemical

23

properties, we combined the cell-adhesive photocrosslinkable elastomeric polymer P(nBAco-%-

24

MABP) with the cell-repellent polymer PAA. Fabrication of substrates in P(nBA-co-%MABP)

25

showed different advantages: i) dilution of polymer in organic solvent allows easy replica-molding

26

fabrication processes; ii) crosslinking the pre-polymer through UV light exposure allow tunable

27

mechanical properties; iii) the polymer is transparent and has a low fluorescence background; iv)

28

crosslinked polymer can be easily functionalized by a layer of proteins that are adsorbed by the

29

surface; v) thanks to the presence of the benzophenone groups, the surface of the polymer can be

30

activated by UV light and easily functionalized.

31

These peculiar characteristics allowed fine tuning of substrate geometry both in 3D (configuration

32

of micro-pillars: cross-section, interspace or height) and in 2D (surface chemistry leading to cell-

33

adhesive and cell-repellent areas). Moreover, the photolithographic process allowed to accurately

34

overlap the surface chemistry pattern with the micro-pillar grid. These specific features allowed for

35

accurate and reproducible large scale hESC-CMs patterning.

20


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Force measurement in cardiac cultures has been matter of several studies in the past, and

2

microstructured substrates found vast application in this field. A number of works employed dense

3

arrays of microposts deflected by contracting cell cultured on top of them [25-28, 32-34, 38]. This

4

experimental approach shares similar downfalls to the more common traction force microscopy

5

with compliant hydrogels embedded with glass or fluorescent beads [35, 50, 51]: it is a 2D system

6

with traction forces derived by tangential shear of the spread cell membrane against a broad

7

surface. The golden standard for force measurement is uni-axial traction measurement and it has

8

been proven to work for engineered cardiac tissues [52] but is not for single cell systems. Only

9

recently, Taylor and colleagues developed the sacrificial layer techniques on which hESC-CMs

10

were able to randomly adhere between two or three adjacent pillars [40]. The authors show a proof

11

of concept that this technique could be used for axial force detection in primary neonatal rat

12

cardiomyocytes by phase contrast imaging.

13

In order to provide a more comprehensive analysis of physiological and pathophysiological

14

mechanisms in heart muscle cells, simultaneous detection of uni-axial force and additional cardiac

15

features (such as calcium handling) would be of great value.

16

For this reason, we developed a biomaterial-based platform that combines the measurement of the

17

axial contractile force of hESCs-CMs with the detection of the corresponding calcium transients

18

using confocal imaging.

19

In order to obtain accurate measurements of forces during contraction, micropillar displacement

20

was correlated to a force value through the application of a finite element model (FEM) developed

21

according to the specific geometry and mechanical properties of the substrate. The contraction

22

force was quantified considering the actual points of cell anchorage to the micropillars, detected by

23

3D reconstruction of cell-pillar system by immunofluorescence for cell focal adhesion and confocal

24

sectioning. This computational effort provided proper analysis of single cell contraction parameters

25

regardless of the individual cell-pillar coupling conditions.

26

We also showed that the morphology and chemistry of the substrate are fitting to long-term cell

27

culture requirements (up to 8 weeks). hESC-CMs on the substrate showed a highly oriented

28

sarcomeric organization. As proof of concept, we applied this experimental setup to follow calcium

29

transients and contraction force evolution of hESC-CMs after 1 week and after 5 weeks of culture

30

on the microstructured substrate. Extension of time in culture has been often reported as a

31

determinant in cardiac structural and functional maturation [9, 48, 49, 53] and we showed here how

32

our experimental setup can be instrumental in analyzing the functional features of human

33

cardiomyocytes maturing in vitro. Cardiomyocytes analyzed after 5 weeks in culture showed on

34

average shorter calcium transients (figure 7B) and an associated increased contraction force (fig

35

7C). Although hPSC-CMs are an early and immature cell type, often referred as fetal-like [8],

36

characterized by absence of structural feature such as t-tubules [54-56], underdeveloped SR [57]

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1

and mixed contribution of SR-stores and trans-sarcolemmal Ca2+ entry to calcium transients [58-

2

60], the hESC-CMs still represent an excellent biological substrate for our experimental platform

3

validation, holding the potential to become a reliable model for the human myocardium in vitro.

4

To our knowledge, this is a first measurement of purely uni-axial contraction force for hESC-CMs,

5

with simultaneous acquisition of calcium dynamics. The feasibility of our experimental approach for

6

contraction force measurements is validated by comparison to the absolute values of reported

7

contraction force measurements made with different experimental setups, although challenging

8

because mostly based on cell in adhesion [41, 42] and reported as tangential stress (i.e. nN/mm2

9

of cell area or nN/post in the case of micro-pillared substrate) [34, 35, 38, 39]. For instances,

10

measurements of 25-45 nN/post have been reported by Rodriguez et al. on human induced

11

pluripotent stem cells-derived cardiomyocytes, using elastomeric force post arrays [38], whereas

12

Ribeiro et al. reported contraction stresses ranging from 0.26 to 0.25 mN/mm2 for both hESC-CM

13

and hiPS-CM [35]. Using micropost arrays, semi-quantitative estimation of 150 nN uni-axial force

14

have been performed by Taylor et al. for neonatal rat cardiomyocytes [40].

15

Altogether, we provide a microengineered platform suitable for precise and high-throughput

16

contraction force measurement in single-cardiomyocyte systems. Pairing force measurements with

17

calcium dynamics increases considerably the experimental value of the acquired data.

18

Furthermore, since micropillar displacement can be simultaneously detected along with other

19

optically encoded measures, further cardiac functional parameters (such as membrane potential

20

detected by fluorescent probes [61]) could be simultaneously detected. With the correct imaging

21

protocol in place, this platform is suitable for high-throughput monitoring of drug-induced changes

22

on more than one cardiac feature at once and can be a valuable tool in drug-screening.

23 24

Conclusions

25

We presented a biomaterial-based platform to perform quantitative in vitro studies on hESC-CMs.

26

In particular we showed simultaneous detection of the uni-axial contraction force and calcium

27

transients of hESCs-CMs with a defined morphology. To guide cell shape, we combined the

28

generation of a 3-dimensional micropillar array with a photochemical surface modification process.

29

As result, an elastomeric micropillar grid coupled with a matrix of single cell adhesive regions were

30

successfully developed. hESC-CMs, selectively grown only on confined areas of the surface,

31

extended between two micropillars and assumed an elongated shape. Calcium transients together

32

with uni-axial contraction force were acquired through confocal analysis. The use of a finite

33

elements model (FEM), which describes the deflection of the micropillars when a tangential force is

34

applied and 3D confocal reconstruction allowed accurate estimation of contraction forces. We

35

reported that the shortening of calcium transient during several weeks of culture is inversely

36

correlated to the uni-axial contraction force detected in single hESC-CMs. The platform can be

22


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suitable for physiological and pathophysiological in vitro assays as well as for monitoring functional

2

maturation of hPSCs-CMs.

3 4

Acknowledgements

5

The authors are grateful for the financial support from the Deutsche Forschungsgemeinschaft

6

(DFG) through the IRTG Soft Matter Science, from the University of Padova through the Progetto

7

Strategico TRANSAC and the grant Giovani Studiosi 2010 (DIRPRGR10), from “Amici del cuore di

8

Montebelluna” and from “ASCM”.

9 10 11 12 13

References

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[1] Yin, S.; Zhang, X.; Zhan, C.; Wu, J.; Xu, J.; Cheung, J. Measuring single cardiac myocyte

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contractile force via moving a magnetic bead. Biophys. J. 2005, 88, 1489-1495.

16

[2] Chang, W.T.; Yu, D.; Lai, Y.C.; Lin, K.Y.; Liau, I. Characterization of the mechanodynamic

17

response of cardiomyocytes with atomic force microscopy. Anal. Chem. 2013, 85, 1395-1400.

18

[3] Walker, C.A.; Spinale, F.G. The structure and function of the cardiac myocyte: a review of

19

fundamental concepts. J. Thorac. Cardiovasc. Surg. 1999, Aug, 118, 375-382.

20

[4] Bers, D. Cardiac excitation-contraction coupling. Nature 2002, 415, 198-205.

21

[5] Koivumaki, J.T.; Takalo, J.; Korhonen, T.; Tavi, P.; Weckstrom, M. Modelling sarcoplasmic

22

reticulum calcium ATPase and its regulation in cardiac myocytes. Philos. Trans. A Math. Phys.

23

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2 3

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1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60 ACS Paragon Plus Environment

Langmuir

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Langmuir

A

A.2 800

Fluo-4 (F/F0)

Ca2+ Force

Contraction force (nN)

1

Week 5

1

800 Ca2+ Force

700 600 500 400 300 200

Fluo-4 (F/F0)

Week 1

0.2

0.4

0.6

0.8

Time (s)

700 600 500 400 300 200

100 0

Contraction force (nN)

A.1

100

0 1.0 1.2

0

0.2

0.4

0.6

0.8

Time (s)

0 1.0 1.2

B

C B.2 Calcium Release

0.75

Week 1 Week 5 0.5

0.25

Fluo-4 (F/F0)

Calcium Reuptake

1000

Week 1 Week 5 0.5

0.25

0

0

E

D

E.1 900

4c

800 700 600 500 400 300 200 0.4

0.6

0.8

1.0

Calcium transient (s)

1.2

Calcium transient (s)

1.4 Week 1 Week 5

1.2 1.0 0.8 0.6 0.4 0.2

Week 1 Week 5

0 Paragon Plus Environment ACS 0.5 0.6 0.7 0.8 0.9 1.0 1.1 1.2 1.3

Beating frequency (Hz)

800

Week 1 Week 5

600 400 200 0

E.2 Max contraction force (nN)

Fluo-4 (F/F0)

0.75

Max contraction force (nN)

B.1

Max contraction force (nN)

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47

900 800 700 600 500 400 300 200

Week 1 100 Week 5 0 0.5 0.6 0.7 0.8 0.9 1.0 1.1 1.2 1.3

Beating frequency (Hz)

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Langmuir

A

Ca2+ Up

Ca2+ Rel

1.5

2.5 Time (s)

Time (s)

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47

1 0.5

2 1.5 1 0.5

0

n=15

n=13

n=16

1

2.5

5

0

n=15

n=13

n=16

1

2.5

5

Fluo-4 concentration (µM)

Fluo-4 concentration (µM)

Ca2+ Rel

Ca2+ Up

B 2.5

**

1

Time (s)

Time (s)

1.5

*

0.5

*

2 1.5 1 0.5

0

n=25

1

n=25

n=27

n=26

n=22

n=25

0 5 10 20 ACS 40Paragon Plus Environment 1

Fluo-4 incubation time (min)

n=25

n=27

n=26

n=22

5 10 20 40

Fluo-4 incubation time (min)

Analysis of Calcium Transients and Uniaxial Contraction Force in

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