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Optics Express

Optics Express

  • Editor: C. Martijn de Sterke
  • Vol. 16, Iss. 12 — Jun. 9, 2008
  • pp: 8594–8603
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A comparative study of living cell micromechanical properties by oscillatory optical tweezers

Ming-Tzo Wei, Angela Zaorski, Huseyin C. Yalcin, Jing Wang, Melissa Hallow, Samir N. Ghadiali, Arthur Chiou, and H. Daniel Ou-Yang  »View Author Affiliations


Optics Express, Vol. 16, Issue 12, pp. 8594-8603 (2008)
http://dx.doi.org/10.1364/OE.16.008594


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Abstract

Micromechanical properties of biological cells are crucial for cells functions. Despite extensive study by a variety of approaches, an understanding of the subject remains elusive. We conducted a comparative study of the micromechanical properties of cultured alveolar epithelial cells with an oscillatory optical tweezer-based cytorheometer. In this study, the frequency-dependent viscoelasticity of these cells was measured by optical trapping and forced oscillation of either a submicron endogenous intracellular organelle (intra-cellular) or a 1.5µm silica bead attached to the cytoskeleton through trans-membrane integrin receptors (extra-cellular). Both the storage modulus and the magnitude of the complex shear modulus followed weak power-law dependence with frequency. These data are comparable to data obtained by other measurement techniques. The exponents of power-law dependence of the data from the intra- and extra-cellular measurements are similar; however, the differences in the magnitudes of the moduli from the two measurements are statistically significant.

© 2008 Optical Society of America

1. Introduction

Mechanical stresses on biological cells have been found to change not only the cell’s morphology, but also cell cycle, gene expression and protein production in many different systems [1–7

1. M. A. Gimbrone Jr., T. Nagel, and J. N. Topper, “Biomechanical activation: an emerging paradigm in endothelial adhesion biology,” J. Clin. Invest. 99, 1809 (1997). [CrossRef]

]. Like many mechanical systems, the stress-strain relationship of biological cells depends on the mechanical integrity of the cells and the extra-cellular matrix around the cells. The mechanical properties of living cells characterized by the frequency dependence of the storage and loss modules were reported by Fabry et al., [8

8. B. Fabry, G. N. Maksym, J. P. Butler, M. Glogauer, D. Navajas, and J. J. Fredberg, “Scaling the microrheology of living cells,” Phys. Rev. Lett. 87, 148102 (2001). [CrossRef] [PubMed]

] who used magnetic tweezers to manipulate ferromagnetic beads attached to the cell membrane. The frequency dependence of the storage modulus (G′) was found to follow a weak power law behavior and was characterized as having soft glassy material properties. In addition, Hoffman et al. [9

9. B. D. Hoffman, G. Massiera, K. M. Van Citters, and J. C. Crocker, “The consensus mechanics of cultured mammalian cells,” Proc. Natl. Acad. Sci. USA 103, 10259 (2006). [CrossRef] [PubMed]

] utilized laser-tracking microrheometry to demonstrate that the complex shear modulus (G*) also follows a weak power law dependence with frequency.

An outstanding question is whether the cell mechanical properties measured by a probe located outside the cell faithfully reflect the intracellular mechanical properties. Due to the lack of consistencies between various existing measurements of the mechanical properties of the cells [10

10. S. Yamada, D. Wirtz, and S. C. Kuo, “Mechanics of Living Cells Measured by Laser Tracking Microrheology,” Biophys. J. 78, 1736 (2000). [CrossRef] [PubMed]

] and the intrinsic inhomogeneity of the cell interior [11

11. M. Mengistu, L. Lowe-Krentz, and H. D. Ou-Yang, “Physical Properties of the Transcytosis Machinery in Endothelial Cells,” Am. Soc. Cell Biology Annual Meeting (2006).

], it is difficult to compare results obtained by different experiments on different cells. To address the question, we set out to conduct experiments that measure cell mechanical properties using probes located exterior to the plasma membrane and naturally occurring probes embedded inside the cell. Specifically, we measured the storage and loss components of the complex shear modulus, G*(ω)=G′(ω)+iG(ω)″, of alveolar epithelial cells using a microrheometer based on oscillatory optical tweezers [12–14

12. L. A. Hough, “Microrheology of Soft Materials Using Oscillating Optical Traps,” Ph.D. thesis, Lehigh University (2003).

]. In this technique, G*(ω) is obtained by trapping and oscillating a submicron endogenous intracellular organelle. Measurements obtained by this technique were compared with results obtained by optical forced oscillation of a 1.5 µm silica bead that was linked to the cytoskeleton through attachment with integrin receptors in the plasma membrane.

Mechanical stresses and the resulting deformations are particularly relevant in lung epithelial cells because they experience a wide variety of mechanical forces in vivo including cyclic stretching, fluid dynamic shear stress, pressure gradients and surface tension forces [7

7. H. C. Yalcin, S. F. Perry, and S. N. Ghadiali, “Influence of airway diameter and cell confluence on epithelial cell injury in an in vitro model of airway reopening,” J. Appl. Physiol. 103, 1796–1807 (2007). [CrossRef] [PubMed]

, 15–18

15. A. M. Bilek, K. C. Dee, and D. P. Gaver, 3rd, “Mechanisms of surface-tension-induced epithelial cell damage in a model of pulmonary airway reopening,” J. Appl. Physiol. 94, 770 (2003).

]. We choose to study lung epithelial cells as a first step to gain a better understanding of pulmonary physiology at the cellular and molecular levels.

2. Experimental setup and methods

2.1 Cell culture

For these studies, human lung epithelial type II cells (CCL-185 from American Type Culture Collection) were cultured in Ham’s F12K cell culture medium containing 10% fetal bovine serum (FBS), antibiotics including 1ml/100ml penicillin/streptomycin (100×) and 1ml/100ml amphotericin B. Cells were seeded onto 30mm diameter cover slips placed in 6 well plates by adding 2ml of a 3×104 cells/ml solution to each well. Cells were grown under standard culture conditions (37°C, 5% CO2 and 95% air) for 24 hours to obtain a sub-confluent monolayer (~25%). Protein-A coated 1.5µm silica beads (G. KISKER GbR, Germany) were treated with the anti-integrin αV antibody (Sigma, CD51) by incubation at 50 µg per 1 mg of beads in 1× phosphate buffered saline (PBS) solution [19

19. D. Choquet, D. P. Felsenfeld, and M. P. Sheetz, “Extracellular matrix rigidity causes strengthening of integrin-cytoskeleton linkages,” Cell 88, 39 (1997). [CrossRef] [PubMed]

]. The bead solution was added to the cell culture well containing the cell-seeded cover slip. After 20 minutes of incubation with the bead solution, the well was washed with PBS solution to remove unbound beads. This process allowed the bead to adhere onto the plasma membrane by the formation of integrin-antibody linkages, as is illustrated in Fig. 1, where the contact area between the cell membrane and the bead is characterized by the bead radius “a” and the half-angle θ.

Fig. 1. A schematic diagram of the integrin anti-body coated beads attaching to the plasma membrane (bead size not to scale).

2.2 Optical tweezer-based microrheometer setup

The use of oscillatory optical tweezers to trap and oscillate a particle embedded in an elastic medium and to measure its mechanical response has been demonstrated elsewhere [12–14

12. L. A. Hough, “Microrheology of Soft Materials Using Oscillating Optical Traps,” Ph.D. thesis, Lehigh University (2003).

, 20

20. L. A. Hough and H. D. Ou-Yang, “Correlated motions of two hydrodynamically coupled particles confined in separate quadratic potential wells,” Phys. Rev. E. 65 (2002). [CrossRef]

]. Fig. 2 shows the schematic diagram of the setup of an oscillating optical tweezer-based microrheometer used in this study. The expanded beam, with a size slightly larger than the back aperture of the microscope objective OBJ to achieve an “overfill, or full numerical aperture” condition. An infrared laser (Nd:YVO4 1064nm diode-pumped solid-state cw laser, Spectra Physics) was steered by a mirror mounted on a PZT piezo-electric 2-axis (orthogonal) tip/tilt platform (S-330, Physik Instrumente) which was connected to a function generator (a built-in function of the lock-in amplifier, Stanford Research SR-830).

The 1064 nm and a 633 nm HeNe laser beams (Uniphase, 5mW) joined with a beam combiner cube into a collinear configuration were launched into the right-side-port of an inverted microscope (Olympus IX-81) via a telescope lens pair, and directed into the direction of microscope optical axis via a dichromatic mirror (not shown). An oil immersion objective lens OBJ (NA=1.3, 100X, Olympus) was used to focus both laser beams to a common focus. The 1064 nm beam was used to trap and the 633 nm beam was used to track the position of dielectric probe particles inside the sample.

Fig. 2. A schematic diagram of the experimental setup. The area enclosed by the dashed lines represents an inverted optical microscope.

The tracking beam, diffracted by the moving probe particle, was collected by the condenser CDSR (NA=0.9) and projected onto a dual-lateral position-sensing detector (PSD: PSM2-4 and OT-301, On-Trak) where the motion of the probe particle in the transverse sample plane (X-Y plane) was tracked in real-time. An optical filter was inserted to prevent 1064 nm beam from reaching the PSD. The lock-in amplifier, fed with signals from PSD and referenced with the sinusoidal electric signal that drove the PZT-driven mirror, provided the displacement amplitude and phase shift of the trapped particle with great sensitivity.

A CCD camera was installed on the microscope left-side-port to record bright-field images. The dichromatic mirror reflected laser beams and passed visible lights from the microscope halogen lamp so that sample imaging and rheological measurements could be performed simultaneously. The lock-in amplifier, PZT-driven mirror, and the microscope (including the CCD camera) were controlled by a PC.

2.3. System calibration

We applied an oscillatory optical force to a trapped bead and analyzed the oscillatory motion of the particle to determine the viscoelastic moduli of the media. The trapped bead was forced to oscillate along the x direction by the oscillatory tweezers driven by the PZT-controlled mirror. For a particle embedded in a viscoelastic medium, the trapped bead experiences a Stokes drag, a force from the oscillatory optical tweezers, and a negligible force produced by the 633 nm beam. The equation of motion of the trapped bead can be written as [12–14

12. L. A. Hough, “Microrheology of Soft Materials Using Oscillating Optical Traps,” Ph.D. thesis, Lehigh University (2003).

]:

mẍ+6πηaẋ+kx=kOT[Acos(ωt)x]
(1)

where m and a are the mass and the radius of the bead, x is the displacement of the bead relative to its unperturbed position, η(ω) and k(ω) are the frequency dependent viscosity and elasticity of the surrounding medium, A and ω are the amplitude and angular frequency of the oscillatory trapping beam and kOT is the spring constant of the optical trap. Taking into account the very low Reynolds number of the system, i.e., 6πηaẋmẍ, we ignore the first term mẍ in Eq. (1) and derive a steady-state solution to Eq. (1) as x(ω, t)=D(ω) cos [ωt- δ(ω)]. The displacement amplitude, D, and phase shift relative to the phase of the oscillating tweezers, δ, can be written as [14

14. M. T. Valentine, L. E. Dewalt, and H. D. Ou-Yang, “Forces on a colloidal particle in a polymer solution: a study using optical tweezers,” J. Phys. Condens. Matter 8, 9477 (1996). [CrossRef]

]

D(ω)=AkOT(kOT+k)2+(6πηaω)2
(2)
δ(ω)=tan1(6πηaωkOT+k)
(3)

The first step in using the tweezer-based rheometer is to determine the trap spring constant kOT by applying the above two equations for water where η(ω) is the viscosity of water, and k(ω) is zero for all frequencies. Experimental data representing the frequency-dependent amplitude D(ω) and phase shift δ(ω) of an oscillating 1.5 µm silica bead in PBS solution and the corresponding theoretical fits are depicted in Fig. 3(a) and 3(b). The spring constant kOT deduced from the best fit is 14.52 pN/µm from the amplitude data and 13.63 pN/µm from the phase data; in this case η is taken to be 0.01 poise, the viscosity of water. We also tracked the Brownian motion of the same silica bead trapped by a stationary trapping beam and analyzed its position distribution and the optical force field using Boltzmann statistics for a particle in a parabolic potential well [21–23

21. M. T. Wei and A. Chiou, “Three-dimensional tracking of Brownian motion of a particle trapped in optical tweezers with a pair of orthogonal tracking beams and the determination of the associated optical force constants,” Opt. Express 13, 5798 (2005). [CrossRef] [PubMed]

]; the optical spring constant deduced from this experiment (data not shown) is 13.85 pN/µm. The average of the two kOT values from the oscillatory tweezers approach (14.08 pN/µm) was used for the extracellular measurements where the same 1.5 µm silica beads were used.

Fig. 3. (a) The normalized amplitude (left), and (b) the relative phase (right) vs. the oscillation frequency of a 1.5 µm diameter silica bead in an oscillatory optical tweezers. The solid dots represent the experimental data and the solid lines are the fits with the spring constant kOT as the only fitting parameter.

3. Experimental Results

3.1. Intracellular Measurements

For intra-cellular probing, we used optical tweezers to trap and oscillate an internal granule (probably a lamellar body which abundantly exist in alveolar epithelial type II cells) [25

25. K. A. Foster, C.G. Oster, M. M. Mayer, M. L. Avery, and K. L. Audus, “Characterization of the A549 cell line as a type II pulmonary epithelial cell model for drug metabolism,” Exp. Cell Res. 243, 359 (1998). [CrossRef] [PubMed]

] in a configuration shown in Fig. 4(A). Fig. 4(B) shows a bright-field image of these lamellar bodies. To further characterize these lamellar bodies, we used an image analysis software (Metamorph) to determine the area (A), perimeter (P) and width (W) of each lamellar body. This data was specifically obtained by using Metamorph’s “Integrated Morphometry Analysis” routine which automatically outlines the lamellar bodies based on appropriate thresholding of the image. Once A, P and W were known, we calculated the diameter as D=4*A/P and a sphericity parameter as S=D/W. Results indicate that the mean diameter for n=24 lamellar bodies was D=1.95 µm with a 95% confidence interval of [1.87µm, 2.08µm]. In addition, the mean sphericity parameter was S=1.05 with a 95% confidence interval of [1.03, 1.09]. Therefore, we conclude that the intracellular granules used in this study (i.e. lamellar bodies) were spherical, had a uniform spatial distribution and were ~2µm in diameter.

Fig. 4. (A) A sketch of optical tweezer-based cytorheometer. Optical tweezers were used to manipulate an intracellular granular structure (lamellar body, left circle), or an extracellular anti-body coated glass bead (right circle). (B) A bright-field image of lamellar bodies that abundantly exist in alveolar epithelial type II cells.

In this case, G′ and G″ were determined from the experimentally measured particle displacement magnitude (D) and phase shift (δ) via the following two equations [12

12. L. A. Hough, “Microrheology of Soft Materials Using Oscillating Optical Traps,” Ph.D. thesis, Lehigh University (2003).

, 13

13. L. A. Hough and H. D. Ou-Yang, “Viscoelasticity of aqueous telechelic poly(ethylene oxide) solutions: Relaxation and structure,” Phys. Rev. E. 73 (2006). [CrossRef]

]:

G(ω)=k(ω)6πa=kOT6πa(Acosδ(ω)D(ω)1)
(4)
G(ω)=ωη(ω)=kOT6πa(Asinδ(ω)D(ω))
(5)

Since these organelles have different optical properties from that of the silica beads, the kOT used for calculating G* must be different from the value described in Sec. 2.3. Because the radiation gradient force scales as the optical contrast: nm(m 2-1)/(m 2+2) [26

26. A. Ashkin, J. M. Dziedzic, J. E. Bjorkholm, and S. Chu, “Observation of a single-beam gradient force optical trap for dielectric particles,” Opt. Lett. 11, 288–290 (1986). [CrossRef] [PubMed]

, 27

27. H. D. Ou-Yang, Polymer-Colloid Interactions: From Fundamentals to Practice (John Wiley and Sons, New York, 1999), Chap. 15.

], where nm is the index of refraction of the surrounding medium, and m is the ratio of refractive index of the particle to the medium, we rescale kOT by taking the ratio of the estimated optical contrast of the organelles [composed of 80% proteins/lipids (n~1.5) and 20% water (n=1.33)] in the cytoplasmic fluids [80% water and 20% proteins] to the optical contract of the silica bead (n=1.45) in water. The rescaling reduced the kOT to approximately 61% of the value determined from a 1.5 µm silica bead in water. We did not consider further scaling of kOT due to the particle size difference between the lamella bodies and the silica particle because the gradient force is a weak function of the particle size for particles larger than the width of the optical trap (~0.8µm) [28

28. A. Rohrbach , “Stiffness of optical traps: quantitative agreement between experiment and electromagnetic theory,” Phys. Rev. Lett. 95, 168102 (2005). [CrossRef] [PubMed]

]. Note all experiments in this study were performed at room temperature.

Experimental data of the viscoelastic moduli obtained by trapping and oscillating a sub-micron intracellular organelle are shown in Fig. 5. Both G′ and the magnitude of the complex modulus |G*| followed a power law relationship with frequency as shown by solid lines in Fig. 5, i.e. G′=G′0(ω/ω0)α and |G*|=G*0(ω/ω0)β. Data similar to those shown in Figure 5 were obtained in n=5 cells. We found that the intracellular G* measurements contained significant fluctuations and G′<G″ at most frequencies. From the measurements in 5 cells the mean value of the power-law exponents were α=0.24±0.09 and β=0.28±0.05, in agreement with Hoffman et al. [9

9. B. D. Hoffman, G. Massiera, K. M. Van Citters, and J. C. Crocker, “The consensus mechanics of cultured mammalian cells,” Proc. Natl. Acad. Sci. USA 103, 10259 (2006). [CrossRef] [PubMed]

] for the intracellular measurements (β=0.26). Here we report the power-law exponents (α and β) as mean ± standard deviation and have used a one-sample Kolmogorov-Smirnov test to demonstrate that the α and β values follow a normal distribution (p=0.94 and p=0.99 for α and β respectively). The scaling parameters G′0 and G*0 were found to follow a log-normal distribution. Specifically, a one-sample Kolmogorov-Smirnov test on the log’ed G′0 and G*0 values indicate that this data is log-normally distributed (p=0.802 and p=0.718 for G′0 and G*0 respectively). Therefore, “log- transformed” statistics were used to calculate the mean and 95% confidence intervals. The mean value for the scaling parameter G0 at a given frequency of 10Hz (ω0=62.8 rad/s) was G′0=181 dyne/cm2, 95% confidence interval [60.5 dyne/cm2, 543 dyn/cm2] and the mean value for G*0 at 10Hz was G*0=221 dyne/cm2, [57.9 dyne/cm2, 845 dyne/cm2].

Fig. 5. Experimental data obtained by using intracellular organelles as probes: (A) G′(ω) and G″(ω), (B) G*(ω). Solid lines represent power-law fits to G′ and G*

3.2. Extracellular Measurements

To apply an extra-cellular probe, we trapped an anti-integrin αV conjugated silica bead (diameter=1.5µm) externally attached to the plasma membrane of the lung epithelial cell (shown in Fig. 1) and oscillated the trapping beam with an amplitude of A=100 nm. We determined the viscoelasticity of the cell from the amplitude (D) and the phase shift (δ) of the oscillating bead by modifying the approach developed for a bead embedded in a homogeneous and isotropic medium (see Eqs. 4 and 5) with an approach reported by Laurent, et al. [24

24. V. M. Laurent, S. Henon, E. Planus, R. Fodil, M. Balland, D. Isabey, and F. Gallet “Assessment of mechanical properties of adherent living cells by bead micromanipulation: comparison of magnetic twisting cytometry vs. optical tweezers,” J. Biomech. Eng. 124, 408 (2002). [CrossRef] [PubMed]

] which accounts for a system in which the bead under a static force is located at the interface of an aqueous medium and a viscoelastic medium. In the case here, the storage modulus G′(ω) and the loss modulus G″(ω) of the viscoelastic medium are given by:

G(ω)=kOT4πa(32sinθ+cosθsin3θ)(Acosδ(ω)D(ω)1)
(6)
G(ω)=3kOT16a sinθ(Asinδ(ω)D(ω))
(7)

Due to the lack of an accurate measurement for the subtending half-angle θ (see Fig. 1), we assume θ=π/4 in all of our calculations. Note that the choice of θ only influences the magnitude of G′ and G″ and does not influence power-law fits (see below). Estimates of G′ and G″ using different values of θ yield a constant multiplicative factor to the values shown in Fig. 6. For θ in the range of 30 – 90 degrees, the multiplicative factors are 2.4 – 0.36 for G′ and 1.4 – 0.71 for G″, respectively. Because this multiplicative factor will be significantly larger when θ is smaller than 30 degrees, we avoided trapping any beads that were only very loosely bound to the cell surface, namely, the beads that visibly moved synchronously with the oscillatory trap at low oscillatory frequencies. By doing so, we also avoided making measurements on beads that were not anchored by integrins. Experimental data obtained for G′ and G″ using an externally attached bead is shown in Fig. 6(A). A power-law fit for the storage modulus (i.e. G′=G′0(ω/ω0)α) is shown as a solid line. The results of |G*| and power-law fit (i.e. |G*|=G*0(ω/ω0)β) are shown in Fig. 6(B). Similar measurements were obtained in n=20 cells and the mean power-law exponents were α=0.19±0.07 and β=0.27±0.06. We have used a one-sample Kolmogorov-Smirnov test to demonstrate that the externally measured α and β values follow a normal distribution (p=0.93 and p=0.80 for α and β respectively) and report α and β values as mean ± standard deviation. We note that these values are in excellent agreement with results obtained by Fabry et al. [8

8. B. Fabry, G. N. Maksym, J. P. Butler, M. Glogauer, D. Navajas, and J. J. Fredberg, “Scaling the microrheology of living cells,” Phys. Rev. Lett. 87, 148102 (2001). [CrossRef] [PubMed]

] (i.e. α=0.17). We also calculated G′0 and G*0 at a frequency of 10Hz (i.e. ω0=62.8 rad/s). Similar to the internal measurements, a one-sample Kolmogorov-Smirnov test on the log’ed G′0 and G*0 values indicate that this data is also log-normally distributed (p=0.99 and p=0.97 for G′0 and G*0 respectively) and “log-transformed” statistics were used to calculate the mean and 95% confidence interval. For external measurements the mean values were G′0=4570 dyne/cm2, [3050 dyne/cm2, 6830 dyne/cm2] and G*0=5080 dyne/cm2, [3410 dyne/cm2, 7570 dyne/cm2]. The mean value of G*0 agrees well with other optical tweezer based measurements on the A549 cell line [29

29. M. Balland, N. Desprat, D. Icard, S. Fεrεol, A. Asnacios, J. Browaeys, S. Hεnon, and F. Gallet, “Power laws in microrheology experiments on living cells: Comparative analysis and modeling,” Phys. Rev. E 74, 021911 (2006). [CrossRef]

].

Fig. 6. (A) G′(ω) and G″(ω), (B) G*(ω) probed with anti-integrin conjugated silica beads attached to the plasma membrane. Solid lines represent power-law fit to G′ and G*.

4. Discussions

Fig. 7. (A) Power-law exponents of G′ and G* for extracellular and intracellular data. Error bars represent standard deviation and means are not statistically different. (B) Magnitudes of prefactor G′o and G*o for extracellular and intracellular data. Error bars are 95% confidence intervals on log-scale and means are statistically different (p<0.01, log-transformed t-test).
Fig. 8. (A) A bright-field image of cells with a membrane-bound bead. (B) A fluorescent image of actin cytoskeleton (phalloidin) with bead location outlined in blue. Images indicate minimal local restructuring of the actin cytoskeleton near the bead due to short duration of incubation (20 min).

5. Conclusion

We used oscillatory optical tweezers to determine the intracellular viscoelasticity of human lung epithelial cells by forced oscillation of (1) an anti-integrin αV conjugated silica bead externally attached to the plasma membrane and, (2) an endogenous intracellular granule. The two methods yielded comparable results. Specifically, the storage modulus G′ and the magnitude of the complex shear modulus |G*| were found to follow weak power law dependence with the oscillation frequency in the range of f=1 Hz~1,000 Hz, which is consistent with the soft glassy behavior of cellular materials. We have demonstrated that the optical forced oscillation of external probes attached to the cytoskeleton can provide reliable measurements of “whole-cell mechanics”, however, in comparison, the intracellular oscillation of endogenous organelles may be more useful in investigating intracellular heterogeneity and temporal fluctuations.

Acknowledgments

MTW and AC were supported by the National Science Council of the Republic of China Grants NSC 95-2752-E010-001-PAE, NSC 94-2120-M-010-002, NSC 94-2627-B-010-004, NSC 94-2120-M-007 -006, and 95A-C-D01-PPG-01. SNG was supported by the American Heart Association Beginning Grant-In-Aid, an NSF CAREER award and is a Parker B. Francis Fellow of Pulmonary Research. AZ, JW and HDO were supported by the National Science Foundation grants: DMR-041259, EEC- 0343283, and the Lehigh Center for Optical Technologies.

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L. A. Hough and H. D. Ou-Yang, “Correlated motions of two hydrodynamically coupled particles confined in separate quadratic potential wells,” Phys. Rev. E. 65 (2002). [CrossRef]

21.

M. T. Wei and A. Chiou, “Three-dimensional tracking of Brownian motion of a particle trapped in optical tweezers with a pair of orthogonal tracking beams and the determination of the associated optical force constants,” Opt. Express 13, 5798 (2005). [CrossRef] [PubMed]

22.

M. T. Wei, K. Yang, A. Karmenyan, and A. Chiou , “Three-dimensional optical force field on a Chinese hamster ovary cell in a fiber-optical dual-beam trap,” Opt. Express 14, 3056 (2006). [CrossRef] [PubMed]

23.

E. L. Florin, A. Pralle, E. H. K. Stelzer, and J. K. H. Horber, “Photonic force microscope calibration by thermal noise analysis,” Appl. Phys A-Mater. 66, S75 (1998). [CrossRef]

24.

V. M. Laurent, S. Henon, E. Planus, R. Fodil, M. Balland, D. Isabey, and F. Gallet “Assessment of mechanical properties of adherent living cells by bead micromanipulation: comparison of magnetic twisting cytometry vs. optical tweezers,” J. Biomech. Eng. 124, 408 (2002). [CrossRef] [PubMed]

25.

K. A. Foster, C.G. Oster, M. M. Mayer, M. L. Avery, and K. L. Audus, “Characterization of the A549 cell line as a type II pulmonary epithelial cell model for drug metabolism,” Exp. Cell Res. 243, 359 (1998). [CrossRef] [PubMed]

26.

A. Ashkin, J. M. Dziedzic, J. E. Bjorkholm, and S. Chu, “Observation of a single-beam gradient force optical trap for dielectric particles,” Opt. Lett. 11, 288–290 (1986). [CrossRef] [PubMed]

27.

H. D. Ou-Yang, Polymer-Colloid Interactions: From Fundamentals to Practice (John Wiley and Sons, New York, 1999), Chap. 15.

28.

A. Rohrbach , “Stiffness of optical traps: quantitative agreement between experiment and electromagnetic theory,” Phys. Rev. Lett. 95, 168102 (2005). [CrossRef] [PubMed]

29.

M. Balland, N. Desprat, D. Icard, S. Fεrεol, A. Asnacios, J. Browaeys, S. Hεnon, and F. Gallet, “Power laws in microrheology experiments on living cells: Comparative analysis and modeling,” Phys. Rev. E 74, 021911 (2006). [CrossRef]

30.

X. Trepat, M. Grabulosa, F. Puig, G.N. Maksym, D. Navajas, and R. Farre, “Viscoelasticity of human alveolar epithelial cells subjected to stretch,” Am. J. Physiol. Lung Cell Mol. Physiol. (2004) July 9. [CrossRef] [PubMed]

OCIS Codes
(140.7010) Lasers and laser optics : Laser trapping
(170.1420) Medical optics and biotechnology : Biology
(170.1530) Medical optics and biotechnology : Cell analysis
(170.4520) Medical optics and biotechnology : Optical confinement and manipulation
(350.4855) Other areas of optics : Optical tweezers or optical manipulation

ToC Category:
Medical Optics and Biotechnology

History
Original Manuscript: January 3, 2008
Revised Manuscript: May 19, 2008
Manuscript Accepted: May 22, 2008
Published: May 28, 2008

Virtual Issues
Vol. 3, Iss. 7 Virtual Journal for Biomedical Optics

Citation
Ming-Tzo Wei, Angela Zaorski, Huseyin C. Yalcin, Jing Wang, Melissa Hallow, Samir N. Ghadiali, Arthur Chiou, and H. Daniel Ou-Yang, "A comparative study of living cell micromechanical properties by oscillatory optical tweezers," Opt. Express 16, 8594-8603 (2008)
http://www.opticsinfobase.org/oe/abstract.cfm?URI=oe-16-12-8594


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References

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  21. M. T. Wei and A. Chiou, "Three-dimensional tracking of Brownian motion of a particle trapped in optical tweezers with a pair of orthogonal tracking beams and the determination of the associated optical force constants," Opt. Express 13, 5798 (2005). [CrossRef] [PubMed]
  22. M. T. Wei, K. Yang, A. Karmenyan, and A. Chiou, "Three-dimensional optical force field on a Chinese hamster ovary cell in a fiber-optical dual-beam trap," Opt. Express 14, 3056 (2006). [CrossRef] [PubMed]
  23. E. L. Florin, A. Pralle, E. H. K. Stelzer, and J. K. H. Horber, "Photonic force microscope calibration by thermal noise analysis," Appl.Phys. A-Mater. 66, S75 (1998). [CrossRef]
  24. V. M. Laurent, S. Henon, E. Planus, R. Fodil, M. Balland, D. Isabey, and F. Gallet "Assessment of mechanical properties of adherent living cells by bead micromanipulation: comparison of magnetic twisting cytometry vs. optical tweezers," J. Biomech. Eng. 124, 408 (2002). [CrossRef] [PubMed]
  25. K. A. Foster, C.G. Oster, M. M. Mayer, M. L. Avery, and K. L. Audus, "Characterization of the A549 cell line as a type II pulmonary epithelial cell model for drug metabolism," Exp. Cell Res. 243, 359 (1998). [CrossRef] [PubMed]
  26. A. Ashkin, J. M. Dziedzic, J. E. Bjorkholm, and S. Chu, "Observation of a single-beam gradient force optical trap for dielectric particles," Opt. Lett. 11, 288-290 (1986). [CrossRef] [PubMed]
  27. H. D. Ou-Yang, Polymer-Colloid Interactions: From Fundamentals to Practice (John Wiley and Sons, New York, 1999), Chap. 15.
  28. A. Rohrbach, "Stiffness of optical traps: quantitative agreement between experiment and electromagnetic theory," Phys. Rev. Lett. 95, 168102 (2005). [CrossRef] [PubMed]
  29. M. Balland, N. Desprat, D. Icard, S. Féréol, A. Asnacios, J. Browaeys, S. Hénon, and F. Gallet, "Power laws in microrheology experiments on living cells: Comparative analysis and modeling," Phys. Rev. E 74, 021911 (2006) [CrossRef]
  30. X. Trepat, M. Grabulosa, F. Puig, G.N. Maksym, D. Navajas, and R. Farre, "Viscoelasticity of human alveolar epithelial cells subjected to stretch," Am. J. Physiol. Lung Cell Mol. Physiol. (2004) July9. [CrossRef] [PubMed]

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