Highly durable crack sensor integrated with silicone rubber cantilever for measuring cardiac contractility

To date, numerous biosensing platforms have been developed for assessing drug-induced cardiac toxicity by measuring the change in contractile force of cardiomyocytes. However, these low sensitivity, low-throughput, and time-consuming processes are severely limited in their real-time applications. Here, we propose a cantilever device integrated with a polydimethylsiloxane (PDMS)-encapsulated crack sensor to measure cardiac contractility. The crack sensor is chemically bonded to a PDMS thin layer that allows it to be operated very stably in culture media. The reliability of the proposed crack sensor has been improved dramatically compared to no encapsulation layer. The highly sensitive crack sensor continuously measures the cardiac contractility without changing its gauge factor for up to 26 days (>5 million heartbeats), while changes in contractile force induced by drugs are monitored using the crack sensor-integrated cantilever. Finally, experimental results are compared with those obtained via conventional optical methods to verify the feasibility of building a contraction-based drug-toxicity testing system.

Normalized displacement for f, flat, and g, nano-grooved silicone rubber cantilevers. Error bars are mean ± s.d. (n=5 biologically independent samples). b, Sarcomere length of cardiomyocytes on day 10 and 22 of culture day. Error bars are mean ± s.d., (n = 10)., * p < 0.05. c, Normalized displacement of the cantilever owing to the contraction force of cardiomyocytes on day 10 and 22 of the culture day., ** p < 0.01. d, mRNA expression (MHC 6, α-actinin) in cardiomyocytes on day 10 and 22 of the culture period., ** p < 0.01. e, Change in cTnT expression of the cardiomyocytes before and after drug treatment on day 10 and 22 of the culture period., ** p < 0.01 measures by two-way ANOVA followed by Tukey's honest significant difference test. f, Change in β-actin expression of the cardiomyocytes on day 10 and 22 of the culture day before and after drug treatment., ** p < 0.01 measures by two-way ANOVA followed by Tukey's honest significant difference test. Error bars are mean ± s.d. (n=5 biologically independent samples).       Figure 1 describes the mechanism of crack sensors that are distinctly different from the commercial strain gage sensor. The cured silicone rubber cantilever integrated with the proposed crack sensor was chemically bonded to PDMS by plasma treatment in an oxygen atmosphere for long-term use in conductive culture medium. Since the metal layer can-not chemically bond with PDMS-encapsulation layer even after the plasma treatment, an adhesion layer of Cr/SiO2 was deposited on the Pt layer for appropriate adhesion between the Pt metal layer and the PDMS layer. The proposed crack sensor retains its non-linear characteristics same as the conventional crack sensor even after embedding the Cr, SiO2, and PDMS.
The physical model in Supplementary Figure 1 illustrates the mechanism of the proposed crack sensor. At the initial state, all the lips of the crack are in a closed state. When strain is applied to the crack sensor (State 1), all the lips are instantly opened, and the resistance increases as the electrical pathways are blocked.
Then the lips are partially closed due to the compressive force of the Poisson's ratio (State 2) and partially open the electrical path. Thus, through the repetition of states 1 and 2, the crack sensor exhibits an exponential resistance increase rather than momentarily increase in resistance. Finally, when the applied strain approaching or exceeding 1% all the lips are opened, and the resistance reaches infinity (state 5).
After withdrawing the applied tension, the gap between the cracks closed, and the initial resistance of the crack sensor is completely restored.
Detailed theoretical modeling and the non-linear characteristics of the crack sensor is explained in detail in the supporting information according to our previous study [1,2].
For that sensor due to a technology of producing large unidirectional strain, the free cracks cut the sensor strip through so that the normalized conductance of the sensor vs strain was determined by the probability distribution function (pdf) ( ) of the steps on a crack lip 15 making contacts between the lips. For a free crack we found a Supplementary Equation for ( ) with the only "size" parameter -the strain 0 that corresponds to the crack gap width 0 being about the grain where = 0 and is the proportionality factor to be defined by relating the crack gap width to the Equation (1) gives for the resistance = 1/ as a function of strain the following: Owing to the non-linear characteristics, the crack sensor shows the considerable resistance ratio variation (R / R0) ~ 90,000 at higher strain range (0-1%) and relatively small variation of ~1.7 at the lower strain range (0-0.3%). However, the strain produced by the cultured cardiomyocytes on the cantilever is minimal. In this present study, the maximum strain produced by the cultured cardiomyocytes on the cantilever is ~0.03%, which causes ~100 µm cantilever displacement. In addition, the gauge factor of the crack sensor is ~156 times higher than the commercial piezoresistive sensor at 0.03% strain (Supplementary Figure 9). Therefore, the crack sensor is highly sensitive enough to detect even the smaller variation in the displacement of the cantilever caused by the contractile force of the cultured Supplementary Note. 11. We have fabricated the crack sensor arrays by integrating three cantilevers in a glass body and analyzed the characteristics of each cantilever by applying a 0.03% strain. The initial resistance of the fabricated cantilever array was found to be ~276 ± 8.2 Ω, and the resistance increased to ~288 ± 8.1 Ω after applying 0.3% of strain. As shown in Supplementary Figure