Bioz Ratings For Life-Science Research

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Article Details

Authors

Rodrigo Gómez-Martínez, Alberto M Hernández-Pinto, Marta Duch, Patricia Vázquez, Kirill Zinoviev, Enrique J de la Rosa, Jaume Esteve, Teresa Suárez, José A Plaza

Journal

Nature nanotechnology

PM Id

23812188

DOI

10.1038/nnano.2013.118

Table of Contents

Abstract

Rodrigo GóMez-Martı́Nez1, Alberto M. HernáNdez-Pinto2, Marta Duch1, Patricia VáZquez2, Kirill Zinoviev1, Enrique J. De La Rosa2, Jaume Esteve1, Teresa SuáRez2 And José A. Plaza1*

Methods

Acknowledgements

Author Contributions

Additional Information

Competing Financial Interests

Abstract

The ability to measure pressure changes inside different components of a living cell is important, because it offers an alternative way to study fundamental processes that involve cell deformation1. Most current techniques such as pipette aspiration2, optical interferometry3 or external pressure probes4 use either indirect measurement methods or approaches that can damage the cell membrane. Here we show that a silicon chip small enough to be internalized into a living cell can be used to detect pressure changes inside the cell. The chip, which consists of two membranes separated by a vacuum gap to form a Fabry–Pérot resonator, detects pressure changes that can be quantified from the intensity of the reflected light. Using this chip, we show that extracellular hydrostatic pressure is transmitted into HeLa cells and that these cells can endure hypo-osmotic stress without significantly increasing their intracellular hydrostatic pressure. Scientific interest in the intersection of microand nanotechnologies with biology has focused on providing new tools to study fundamental questions in cell biology5–7. Fabrication based on these techniques offers the potential to develop integrated devices with nanosized moving parts8 and allows for new opportunities for the mechanical analysis of cells1,9,10. However, the use of the devices has been focused on extracellular or invasive techniques11. Microand nanoparticles can be internalized inside living cells and have been used in numerous studies in cell biology. Furthermore, silicon-based particles have revealed their superiority in biological imaging and drug delivery because of their inherent biocompatibility12,13. Recently, for the purposes of single-cell labelling, we demonstrated a technique for the fabrication of silicon microparticles based on semiconductor technologies14,15. Using chemical functionalization, we also proved that these microparticles could react with the intracellular medium16. Existing techniques for the indirect measurement of intracellular pressure include methods that induce a large deformation of the cell by aspiration2, or methods that detect variations in the cell volume1,3. In contrast, the servo-null technique allows for a direct measurement by inserting a micropipette as a pressure probe4, but in this process the cell membrane is mechanically damaged. Thus, the measurement of extracellular loads transmitted to the interior of the cell, and in particular to a subcellular component, has not been demonstrated directly. Indeed, the cell is a highly complex and virtually unexplored mechanical system in which the membranes, cytoskeleton and extracellular matrix provide structural integrity. Here, we have fabricated a nanomechanical chip that can be internalized to detect intracellular pressure changes within living cells, enabling an interrogation method based on confocal laser scanning microscopy (CLSM). The design comprises a mechanical sensor (Fig. 1a) defined by two membranes separated by a vacuum gap, and an optical reference area. The membranes act as parallel reflecting mirrors, constituting a Fabry–Pérot resonator that is partially transparent for some wavelengths17. An external pressure P deflects the membranes and changes the gap, tgap (Fig. 1b). Accordingly, the intensity of the reflected light at the centre of the membranes, Ir_Centre , for a given wavelength l, is modulated by P. The reference area is used for focusing purposes. Briefly, the sensing principle is based on the acquisition of images for a given l and the quantification of Ir_Centre. The fabrication processes included the deposition of three structural and three sacrificial layers, poly-silicon and silicon oxide, respectively (Fig. 1c, Supplementary Fig. S1). Polycrystalline silicon was selected as the structural material because of its elastic behaviour and high reliability18. The lateral dimensions of the mechanical membranes were fixed to 3 mm × 3 mm (Fig. 1d). Analytical and simulated analyses showed that the mechanical deformation was highly dependent on the membrane thickness and the linear response versus P (Fig. 1e, Supplementary Fig. S2). We therefore selected 50-nm-thick membranes to achieve a theoretical mechanical sensitivity of 5.5 nm per bar. The high refraction index of poly-silicon gives a spectral selectivity to the structure and, subsequently, a high sensitivity to P. Theoretically, the optical reflection of the structure (Fig. 1f) showed a resonance valley that was a function of tgap and l (Fig. 1g,h). Thus, P shifted the reflection curve towards smaller values ( 2 × Dtgap) and, for fixed l, large variations in the reflection could be obtained. Finally, tgap ≈ 300 nm was selected by considering the high optical sensitivity and cell internalization capabilities. The fabricated devices were validated using a bright-field optical microscope. The experiment showed a minimum reflection for l≈ 570 nm (Fig. 2a). For fixed l, Ir_Centre increased versus P for l . 580 nm, and decreased for l , 560 nm. CLSM images with superior resolution allowed an image-processing algorithm to be developed to detect the pressure loads based on a quantification of the mean intensities of three regions of interest (Supplementary Figs S3–S5). External pressure was applied from 0 to 1 bar and from 1 to 0 bar. Ir_Sensor decreased with a laser wavelength of 514 nm and increased with 594 nm (Fig. 2b). To test the sensor inside living cells, we took advantage of our previous experience of internalizing silicon microparticles inside HeLa cells by lipofection16. Sensors were easily localized by optical light microscopy because of the higher reflectivity of the poly-silicon, and CLSM showed the specific location of the chip in the cytoplasm (Fig. 3a,b). The internalized sensors only represented 0.2% of the total volume of a typical HeLa cell (Supplementary Fig. S6). After transfection, a number of HeLa cells in the culture displayed vacuoles as a result of the lipofection procedure. Our experiments showed that these vacuoles did not affect cell fitness or viability (Fig. 3a,b,

Silicon chips detect intracellular pressure changes in living cells

Rodrigo Gómez-Martı́nez1, Alberto M. Hernández-Pinto2, Marta Duch1, Patricia Vázquez2, Kirill Zinoviev1, Enrique J. de la Rosa2, Jaume Esteve1, Teresa Suárez2 and José A. Plaza1*

The ability to measure pressure changes inside different components of a living cell is important, because it offers an alternative way to study fundamental processes that involve cell deformation1. Most current techniques such as pipette aspiration2, optical interferometry3 or external pressure probes4 use either indirect measurement methods or approaches that can damage the cell membrane. Here we show that a silicon chip small enough to be internalized into a living cell can be used to detect pressure changes inside the cell. The chip, which consists of two membranes separated by a vacuum gap to form a Fabry–Pérot resonator, detects pressure changes that can be quantified from the intensity of the reflected light. Using this chip, we show that extracellular hydrostatic pressure is transmitted into HeLa cells and that these cells can endure hypo-osmotic stress without significantly increasing their intracellular hydrostatic pressure. Scientific interest in the intersection of micro- and nanotechnologies with biology has focused on providing new tools to study fundamental questions in cell biology5–7. Fabrication based on these techniques offers the potential to develop integrated devices with nanosized moving parts8 and allows for new opportunities for the mechanical analysis of cells1,9,10. However, the use of the devices has been focused on extracellular or invasive techniques11. Microand nanoparticles can be internalized inside living cells and have been used in numerous studies in cell biology. Furthermore, silicon-based particles have revealed their superiority in biological imaging and drug delivery because of their inherent biocompatibility12,13. Recently, for the purposes of single-cell labelling, we demonstrated a technique for the fabrication of silicon microparticles based on semiconductor technologies14,15. Using chemical functionalization, we also proved that these microparticles could react with the intracellular medium16. Existing techniques for the indirect measurement of intracellular pressure include methods that induce a large deformation of the cell by aspiration2, or methods that detect variations in the cell volume1,3. In contrast, the servo-null technique allows for a direct measurement by inserting a micropipette as a pressure probe4, but in this process the cell membrane is mechanically damaged. Thus, the measurement of extracellular loads transmitted to the interior of the cell, and in particular to a subcellular component, has not been demonstrated directly. Indeed, the cell is a highly complex and virtually unexplored mechanical system in which the membranes, cytoskeleton and extracellular matrix provide structural integrity. Here, we have fabricated a nanomechanical chip that can be internalized to detect intracellular pressure changes within living cells, enabling an interrogation method based on confocal laser scanning microscopy (CLSM). The design comprises a mechanical sensor (Fig. 1a) defined by two membranes separated by a vacuum gap, and an optical reference area. The membranes act as parallel reflecting mirrors, constituting a Fabry–Pérot resonator that is partially transparent for some wavelengths17. An external pressure P deflects the membranes and changes the gap, tgap (Fig. 1b). Accordingly, the intensity of the reflected light at the centre of the membranes, Ir_Centre , for a given wavelength l, is modulated by P. The reference area is used for focusing purposes. Briefly, the sensing principle is based on the acquisition of images for a given l and the quantification of Ir_Centre. The fabrication processes included the deposition of three structural and three sacrificial layers, poly-silicon and silicon oxide, respectively (Fig. 1c, Supplementary Fig. S1). Polycrystalline silicon was selected as the structural material because of its elastic behaviour and high reliability18. The lateral dimensions of the mechanical membranes were fixed to 3 mm × 3 mm (Fig. 1d). Analytical and simulated analyses showed that the mechanical deformation was highly dependent on the membrane thickness and the linear response versus P (Fig. 1e, Supplementary Fig. S2). We therefore selected 50-nm-thick membranes to achieve a theoretical mechanical sensitivity of 5.5 nm per bar. The high refraction index of poly-silicon gives a spectral selectivity to the structure and, subsequently, a high sensitivity to P. Theoretically, the optical reflection of the structure (Fig. 1f) showed a resonance valley that was a function of tgap and l (Fig. 1g,h). Thus, P shifted the reflection curve towards smaller values ( 2 × Dtgap) and, for fixed l, large variations in the reflection could be obtained. Finally, tgap ≈ 300 nm was selected by considering the high optical sensitivity and cell internalization capabilities. The fabricated devices were validated using a bright-field optical microscope. The experiment showed a minimum reflection for l≈ 570 nm (Fig. 2a). For fixed l, Ir_Centre increased versus P for l . 580 nm, and decreased for l , 560 nm. CLSM images with superior resolution allowed an image-processing algorithm to be developed to detect the pressure loads based on a quantification of the mean intensities of three regions of interest (Supplementary Figs S3–S5). External pressure was applied from 0 to 1 bar and from 1 to 0 bar. Ir_Sensor decreased with a laser wavelength of 514 nm and increased with 594 nm (Fig. 2b). To test the sensor inside living cells, we took advantage of our previous experience of internalizing silicon microparticles inside HeLa cells by lipofection16. Sensors were easily localized by optical light microscopy because of the higher reflectivity of the poly-silicon, and CLSM showed the specific location of the chip in the cytoplasm (Fig. 3a,b). The internalized sensors only represented 0.2% of the total volume of a typical HeLa cell (Supplementary Fig. S6). After transfection, a number of HeLa cells in the culture displayed vacuoles as a result of the lipofection procedure. Our experiments showed that these vacuoles did not affect cell fitness or viability (Fig. 3a,b, 1Instituto de Microelectrónica de Barcelona, IMB-CNM (CSIC), Esfera UAB, Campus UAB, 08193, Cerdanyola, Barcelona, Spain, 2Centro de Investigaciones Biológicas, CIB (CSIC), C/Ramiro de Maeztu 9, 28040, Madrid, Spain. *e-mail: joseantonio.plaza@imb-cnm.csic.es LETTERS PUBLISHED ONLINE: 30 JUNE 2013 | DOI: 10.1038/NNANO.2013.118 NATURE NANOTECHNOLOGY | VOL 8 | JULY 2013 | www.nature.com/naturenanotechnology 517 © 2013 Macmillan Publishers Limited. All rights reserved Supplementary Movie S1) and disappeared when cells were returned to normal culture conditions (Supplementary Fig. S7). Sensorcontaining HeLa cells, with or without vacuoles, divided normally (Fig. 3c, Supplementary Movie S2), displayed active mitochondria (Supplementary Fig. S7) and were healthy 9 days later (Supplementary Fig. S8). We also confirmed that the pH of the vacuoles oscillated between 4 and 6, and that the internalized devices were not degraded inside the HeLa cells 9 days after lipofection (Supplementary Fig. S8). This result is in good agreement with the finding that poly-silicon did not degrade in solutions buffered at pH values between 4 and 9 (Supplementary Fig. S8). We next analysed the mechanical transmission of extracellular pressure to a subcellular component. The presence of a sensor inside a vacuole has several inherent advantages. First, it can give information about how an external pressure is transmitted mechanically to organelles. Second, it prevents the eventual existence of mechanical cross-sensitivity on the devices because of other organelles or cytoskeletal filaments, which can induce small forces and displacements (Supplementary Fig. S2). Third, better-quality CLSM images are obtained when the sensors are immersed in a medium with a uniform refractive index (Supplementary Fig. S9). Figure 4a presents overlaid images of transmitted light and laser channels in which the vacuole and different parts of the device can be easily recognized. An external pressure was applied from 0 to 1 bar and from 1 to 0 bar. A comparison between Ir_Sensor inside the vacuole and for the calibrated sensor in air showed close proportional changes (Figs 4b and 2b) and confirmed that the extracellular pressure is transmitted into the vacuole (Supplementary Fig. S10). The results in Fig. 4c demonstrate the capability of detecting pressure fluctuations inside a cell. The reflection from the sensor depends on the optical properties of the surrounding media; however, the position of the resonance is almost invariant (Supplementary Fig. S11). We also observed that Ir_Sensor is reversible, demonstrating that the pressure inside the Mechanical pressure sensorOptical reference area 4 μm 6 μm 400 nm P = 0 P > 0 P P I(λ) I(λ) I(λ) Ir_Ref (λ) Ir_Ref (λ) Ir_Centre(λ, P = 0) Ir_Centre(λ, P > 0) I(λ) λ (nm) 0 5 10 15 20 25 40 50 60 70 80 90 100 D is pl m em b ( nm ) tmemb (nm) 0 5 10 15 0.0 0.5 1.0 1.5 2.0 Si substrate Ox2 Ox1 Ox3Poly1 Poly2 Poly3 Poly-silicon Vacuum nVacuum nMedium Poly-silicon nMediumMedium Medium nPoly nPoly tTop memb tBottom memb tgap(P) 0.0 0.2 0.4 0.6 0.8 1.0 150 250 350 450 550 650 750 1.0 0.8 0.6 0.4 0.2 0.0 500 600 700 800 Re fle ct io n Re fle ct io n tgap = 280 nm tgap = 290 nm tgap = 300 nm tgap (nm) tgap(P = 0) tgap(P > 0) D is pl m em b ( nm ) P (bar) tmemb = 50 nm 2.0 bar 1.0 bar 0.5 bar λ = 514 nm λ = 561 nm λ = 594 nm a c d e f g h b NATURE NANOTECHNOLOGY | VOL 8 | JULY 2013 | www.nature.com/naturenanotechnology518 © 2013 Macmillan Publishers Limited. All rights reserved b 0 10,000 20,000 30,000 40,000 50,000 60,000 0.00 0.25 0.50 0.75 1.00 0.75 0.50 0.25 0.00 I r_ Se ns or (a .u .) P (bar) λ = 514 nm I r_ Se ns or (a .u .) P (bar) λ = 561 nm 0 10,000 20,000 30,000 40,000 50,000 60,000 0.00 0.25 0.50 0.75 1.00 0.75 0.50 0.25 0.00 I r_ Se ns or (a .u .) λ = 594 nm P (bar) 0.00 0.25 0.50 0.75 1.00 0.75 0.50 0.25 0.00 0 10,000 20,000 30,000 40,000 50,000 60,000 P = 0 ba r P = 1 b ar 500 nma 510 nm 520 nm 530 nm 540 nm 550 nm 560 nm 570 nm 580 nm 590 nm 600 nm 610 nm 620 nm 630 nm 640 nm 650 nm 0 50 100 150 200 250 500 P = 0 bar P = 1 barI r_ C en tr e/ I r_ Re f ( a. u. ) λ (nm) 510 520 530 540 550 560 570 580 590 600 610 620 630 640 650 Figure 2 | Validation of the sensing principle. a, Bright-field optical microscopy experiment in air medium. Top: experimental true-colour images taken by an eight-bit colour CCD camera versus l and P. Band-pass filters from 500 to 650 nm were used to select the working l. Bottom: normalized Ir_Centre/Ir_Ref (255 a.u. for l¼ 500 nm). P induces a lateral displacement of the curve towards smaller l. For fixed l, positive or negative sensitivities are observed (black arrows). Ovals indicate light colour (l¼ 510 nm, blue; l¼ 590 nm, yellow). Error bars,+10% (based on measurement uncertainty from images). b, CLSM experiment in air medium: Ir_Sensor versus P from 16-bit images. Lasers with l¼ 514 nm, 561 nm and 594 nm were used to select the working l. Positive and negative sensitivities are also observed for l¼ 594 nm and l¼ 514 nm, respectively. Ir_Sensor decreased for 0 ≤ P ≤ 0.75 bar and increased for P¼ 1 bar, l¼ 561 nm, as this is close to the resonance valley of the Fabry–Pérot spectrum. Error bars,+5%, 9% and 8% for l¼ 514 nm, 561 nm and 594 nm, respectively (based on measurement uncertainty from images). confocal images. Bottom: orthogonal projection of confocal images showing that the chip is inside the cell. c, A HeLa cell containing a device inside the vacuole can proceed through mitosis (individual frames taken from Supplementary Movie S2; the time format is hh:mm). Scale bars, 10 mm. NATURE NANOTECHNOLOGY | VOL 8 | JULY 2013 | www.nature.com/naturenanotechnology 519 © 2013 Macmillan Publishers Limited. All rights reserved Visible (T) + 514 nm (R)Visible (T) Visible (T) + 561 nm (R) Visible (T) + 594 nm (R) 0 10,000 20,000 30,000 40,000 50,000 60,000 0.000.250.500.75 1.00 0.750.500.250.00 I r_ Se ns or (a .u .) P (bar) λ = 514 nm I r_ Se ns or ( a .u .) 0.000.25 0.50 0.75 1.00 0.75 0.50 0.250.00 P (bar) 0 10,000 20,000 30,000 40,000 50,000 60,000 λ = 561 nm I r_ Se ns or (a .u .) 0.000.25 0.50 0.75 1.00 0.75 0.50 0.250.00 P (bar) λ = 594 nm 0 10,000 20,000 30,000 40,000 50,000 60,000 P (bar) r ( % ) 0.00 0.25 0.50 0.75 1.00 0 25 50 75 Inside cell In air ** * *** ** ** * 420 440 460 480 500 520 0 10 20 λ (n m ) Estimated λ-shift induced by P = 7 bar Outside vacuole Inside vacuole Assay number Before osmotic shock After osmotic shock a b c d such further developments include thinner mechanical layers, autofocus and tilt-stage systems, and computer-assisted measurements. Mechanical forces are not very well understood and are involved in basic cellular processes such as cell migration25,26, diseases27–29 and development30. Intracellular mechanical sensors will provide information directly from inside the cellular environment about these cellular forces and will provide new opportunities. We believe that this is a first step towards a wide-ranging field of intracellular nanochips that will offer a different perspective on fundamental problems in cell biology.

Methods

Imaging acquisition during pressure experiments. For Bright-field optical microscopy, experiments were performed with an Eclipse ME600 upright optical microscope (Nikon). A ×100 magnification, 0.8 NA, long-distance objective LU Plan ELWD 3.5 (Nikon) was used. Images were recorded using an 8-bit colour CCD (chargecoupled device) camera (DXM1200F, Nikon) using the advanced control software Nikon ACT-1 (Automatic Camera Tamer). Band-pass filters (Thorlabs) coupled with a YM-NCB11 filter slider (Nikon) were used to select the wavelength of the incident light. For CLSM, confocal images were acquired with a confocal Leica TCS-SP5 microscope (Leica Microsystems GmbH), using 514 nm, 561 nm and 594 nm excitation laser wavelengths (acousto-optical tunable filters (AOTF)¼ 1%) for the first batch of fabricated chips, and 458 nm, 476 nm, 488 nm, 496 nm and 514 nm excitation laser wavelengths (AOTF¼ 1%), for the second batch of fabricated chips. The confocal analysis was conducted in the acousto-optical beam splitters (AOBS) reflection mode, with 16 bit-depth resolution and in the X–Y–Z scan mode. A ×63/0.9 HCX APO water objective (Leica Microsystems GmbH) was used. The image acquisition time was 25 s. The images were pre-analysed by LAS AF software (Leica Microsystems GmbH). Cell manipulation and osmotic shock. Chips were lipofected inside human HeLa cells using a protocol we have described previously16. HeLa cells were incubated for 12–16 h in the lipofection medium. Cell viability was analysed by incubating cells with Cell Tracker Green and MitoTracker Red (Molecular Probes, Invitrogen) for 15 min at 37 8C. HeLa cells were fixed with 4% paraformaldehyde in PBS for 45 min. The nuclei were stained with DAPI (Molecular Probes) and the cells were mounted with Fluoromont-G (Southern Biotech) for microscopy. HeLa cells were also incubated with Calcein AM, MitoTracker Red, DiOC and Lysosensor Red (Molecular Probes, Invitrogen), for direct observation under the CLSM following the manufacturers’ recommendations. Cells were grown on glass coverslips and observed under the CLSM inside a live-imaging Ludin chamber. To expose cells to an osmotic shock, standard DMEM medium with 10% fetal bovine serum (Molecular Probes, Invitrogen) was 10% diluted in deionized water and perfused in the Ludin chamber. Cell viability imaging. Cells were observed under a TCS SP2 AOBS CLSM with ×63 oil immersion lens (Leica Microsystems GmbH). Green fluorescence was monitored with excitation and emission settings of 488 nm and 505–550 nm, respectively. Red fluorescence was monitored with excitation and emission settings of 561 nm and 580–610 nm, respectively. A 351 nm laser line was used to image nuclei, and fluorescence emission was measured at 415–460 nm. Chips were imaged with a 488 nm laser line and they were detected by reflected light at 480–495 nm. Time-lapse microscopy was performed with a Leica AF6000 LX model DMI6000B, and pictures were taken every 10 min. HeLa cell videos were processed with Leica imaging software. Statistical analysis. Data analysis was performed with Graph Pad Prism 4 software. Analysis of variance (ANOVA) and a Bonferroni test were used to compare intra-group data (‘chip inside cell’ or ‘chip in air’ data sets), and the x2 test was used to compare pressure data from calibration chips in air versus chips inside cells. Extrapolate l for minimum reflection. Values of l for the minimum reflection (Fig. 4d) were extrapolated from data (Supplementary Fig. S14) by adjusting the mean intensities for the five selected lasers to a second-order polynomial. The minimum corresponded to the l where the first derivative of the function was zero. Received 2 August 2012; accepted 24 May 2013; published online 30 June 2013

Acknowledgements

This work was supported by the Spanish Government grants TEC2009-07687-E, TEC2011-29140-C03-01 and SAF2010-21879-C02-01. P.V. was supported by Centro de Investigación Biomédica en Red de Diabetes y Enfermedades Metabólicas Asociadas– Instituto de Salud Carlos III (CIBERDEM-ISCIII). The authors thank M. Calvo of Centros Cientı́ficos y Tecnológicos–Universidad de Barcelona (CCiT-UB), M.T. Seisdedos (CIB), J. Monteagudo (Leica Microsystems S.L.) and D. Megias of Unidad de Microscopı́a ConfocalCentro Nacional de Investigaciones Oncológicas (CMU-CNIO) for their assistance with CLSM experiments and A. Bosch (CCiT-UB) for assistance with image processing. The authors also thank the cleanroom staff of IMB-CNM for fabrication of the chips.

Author contributions

All authors discussed the results and contributed to writing the manuscript. M.D., R.G-M. and J.E. conceived and guided chip fabrication. Optical design and analysis was carried out by K.Z. The biological experiments were performed by A.M.H.P. and P.V., designed by A.M.H.P. and E.J.d.l.R., and planned and coordinated by T.S. R.G-M. performed the experimental characterization of the chips as well as data analysis. J.A.P. conceived and directed the project.

Additional information

Supplementary information is available in the online version of the paper. Reprints and permissions information is available online at www.nature.com/reprints. Correspondence and requests for materials should be addressed to J.A.P.

Competing financial interests