Patent Application: US-22446807-A

Abstract:
a chemical assay on sensing objects e . g . for use as a drug screening assay on living cells , as well uses and a method for making such an integrated system is proposed . the assay is comprising a base element with on a surface an array of multiple immobilisation points for individual sensing objects such as cells or groups of a few sensing objects , and a flow chamber bordered on a first lateral side by said base element and covering said base element at least in the region with the array of immobilisation points , wherein the flow chamber on an entry - side comprises at least one or two inlets for the introduction of different test solutions into the flow chamber in a flow direction , and on an exit - side located opposite to the entry - side comprises at least one outlet for the test solutions , wherein these inlets are located substantially in a plane parallel to the surface of the base element and spaced apart in a direction perpendicular to the flow direction of the test solutions such that the test solutions flow across over the array of multiple immobilisation points and sensing objects located thereon in a parallel laminar flow , such that there is no interference and / or or well - defined and reproducible interference between the flow of the different test solutions over defined groups of the array of multiple immobilisation points .

Description:
in the following and with reference to the drawings , the invention as generally outlined above shall be illustrated . the now following description is however for the purpose of illustrating the present preferred embodiments of the invention and shall not be construed for the purpose of limiting the same as defined in the appended claims . to achieve low reagent consumption in a highly parallel drug - screening approach with integrated detecting or sensing step , a miniaturized equivalent of a micro titer - plate and a dilution stage , both integrated in one system , are desired , so that several functions such as the immobilization and culturing of cells inside an incubation chamber , the drug dilution , and the drug - screening functions have to be integrated . the immobilization of cells can be achieved using methods such as physical retention chambers , where cells are trapped by an inserted cellulose - nitrate membrane [ 1 ], di - electrophoretic methods using an inhomogeneous electrical field [ 2 , 3 ], or the capturing of single cells either at the entrance of a silicon channel [ 4 ] or by pneumatic anchoring [ 5 , 6 ]. also , multi - height ‘ sandbag ’- type structures have been proposed for particle trapping [ 7 ]. in addition to physical methods , surface - chemical strategies such as the use of adhesion proteins patterned by photolithography [ 8 ], micro - contact printing [ 9 ] or the use of self - assembled mono - layers , are promising approaches to facilitate the immobilization of cells on a chip surface . the mixing of a drug and a buffer solution to produce a wide concentration range is needed for drug - screening experiments . as manual dilution is hard to perform on low volumes , micro - fluidic diluters based on polymeric or inorganic materials have been developed by several groups . serial [ 10 ] and combined serial and parallel mixing [ 11 ], combinatorial 3d mixing over several flow magnitudes [ 12 ] and the use of dilution gradients [ 13 ] have been proposed . as micro - fluidic mixers usually operate in the low - reynold &# 39 ; s - number regime , chaotic mixing has been introduced to improve the mixing of the respective drug and buffer solutions [ 14 , 15 ]. although a variety of miniaturized dilution stages have been reported , the majority is limited to a dilution range of about two orders of magnitude . here , we present a microchip - based system containing a miniaturized equivalent of a micro - titerplate as well as a micro - fluidic dilution cascade ( fig1 ). the device can be used for all essential steps of the screening process : ( i ) immobilization of a defined number of cells to yield a homogeneous array , ( ii ) drug dilution , ( iii ) incubation , and ( iv ) optical interrogation . the core of this system is a 7 × 5 - mm 2 silicon chip 1 with an array of 1000 orifices 5 for cell trapping . as this chip 1 does not provide enough area for the micro - fluidic mixers , those are cast on a 2 × 2 - cm 2 poly ( dimethylsiloxane ) ( pdms ) elastomer substrate 3 . after assembly , the diluter , a 0 . 5 - μl incubation chamber 8 and the cell - loading ports 9 constitute a single unit ( fig2 ). the diluter has two inlets 6 , 7 for the cell medium and the drug stock solution , both of which are subsequently mixed in a cascading channel system ( relative concentrations : 100 %, 10 %, 1 %, 0 . 1 %, 0 % of the original drug stock solution ). additionally , the system features two cell loading ports 9 to load the cells into the incubation chamber 8 and to regularly exchange the medium during pre - screening incubation . the experimental data presented in this paper illustrate that this hybrid microsystem allows for performing a drug - screening assay for 5 sample concentrations with only 0 . 4 μl / min of the sample drug . a schematic of the device is shown in fig1 . the microsystem consists of three distinct components : ( a ) a 7 × 5 - mm 2 silicon chip 1 with an array of 1000 orifices 5 for cell trapping , ( b ) 2 × 2 - cm 2 elastomeric substrate 8 , into which the chip 1 is embedded , to enlarge the real estate of the device , and ( c ) a micro - fluidic cover 3 with the integrated diluter cascade , made of pdms . a cell screening with this device is performed as follows : first , a cell suspension is pumped through the incubation chamber , and the cells 4 are trapped on the orifices 5 . this assures a homogeneous cell distribution inside the chamber . then , the excess cells are washed away by a laminar buffer stream to leave the chamber with a defined number of cells in a homogeneous arrangement . cells are only immobilized during loading and can afterwards proliferate freely during the incubation step . the cells are typically incubated for several days before the actual screening process is performed . for screening , only a minute amount of the drug is pumped into one inlet of the dilution cascade , where it is mixed with a buffer solution from the other inlet to yield the relative final concentrations of 100 %, 10 %, 1 %, 0 . 1 %, 0 % of the drug . the five diluter outputs 17 provide laminar streams over the respective areas of the immobilized cells , so that each stream only perfuses a defined part of the overall cell area . simultaneously or sequentially , the cellular response can be optically assessed by e . g . adding specific fluorescent tags to the buffer stream . the cells 4 are immobilized on the silicon chip i by individual trapping on an array of 5 × 200 orifices 5 owing to a slight pressure difference between the inside and the outside of the incubation chamber 8 . typically , a single cell is immobilized on one orifice 5 during this process . this technique , denoted as ‘ pneumatic anchoring ’, has been previously described by [ 5 ] and by our group [ 6 ] for bio - electronic cmos chips . cell immobilization is used here for mainly two reasons . first , the technique allows for loading the chamber with an exactly defined number of cells for each experiment . consequently , the resulting fluorescence intensity measurements lead to reproducible and statistically relevant data for the different drug concentrations . second , a homogeneous cell carpet is obtained owing to the equal spacing between the orifices ; without immobilization features , the cell loading would lead to irreproducible and spatially imbalanced cell populations that are not suitable for screening experiments . after the loading step has been completed , the immobilization force has been found to be not to disturb the cell proliferation . although cells might migrate during the incubation , the homogeneous nature of the cell carpet is preserved . as the diameter of the cells used in this project , normal human dermal fibroblasts ( nhdfs ), is approximately 20 μm , orifices in the range of 5 μm need to be fabricated to prevent any suction of the cells through the orifices . silicon was used as the chip material , because of the available precision etching techniques . orifices 5 have been etched from the frontside by reactive - ion etching , their back - side has been thinned by anisotropic wet etching to a 5 - μm membrane to reduce the lateral widening of the orifices during fabrication . as silicon technology is comparably expensive , the chip size is limited to the absolutely necessary area ( 7 × 5 mm 2 ). to have enough space around this chip for the integration of the micro - fluidics , the chip has been seamlessly embedded into a larger , 2 × 2 - cm 2 pdms substrate 2 before the micro - fluidic cover 3 has been bonded onto the chip 1 . no leakage of drugs into the cleft between the chip and the micro - fluidic system has been observed . all the necessary parts for the drug handling have been integrated into the micro - fluidic cover ( fig2 ). the orifice array is covered by a 0 . 5 - μl incubation chamber 8 ( 3 . 5 mm wide , 1 . 4 mm long , and 100 μm high ). two loading ports 9 ( 5 mm long ) have been provided to inject the cell suspension into the incubation chamber 8 . the cell loading stream is perpendicular to the main buffer stream . two inlets 6 , 7 are provided for the buffer solution and the drug stock which are mixed in the cascading dilution stage to produce the desired concentrations . five outlets 17 ( 100 μm wide , 700 μm spacing ) provide the drug dilution to five cell arrays . on the opposite side of the chamber , a symmetrical shape port 10 leads to the waste reservoir . as micro - fluidic devices generally operate in the laminar - flow regime , mixing in the dilution stage is only achieved by diffusion . for the structure presented here this also holds true as the reynold &# 39 ; s number is between 0 . 1 and 2 , which is far below the threshold for turbulent flow . to ensure complete mixing , the channel geometries have to be adapted in terms of width and length , and the corresponding flow rates have to be chosen accordingly . the mixing ratios are defined by the flow rates of the drug and the buffer solution at the branches of the diluter stage . at each interception point , the flow rate of the incoming drug ( or output of the previous dilution stage ) is 9 times smaller than the flow rate of the buffer to obtain the desired dilution of 1 : 9 . thus , using three cascading levels with three interception points , dilutions of 10 %, 1 %, 0 . 1 % can be achieved . this modular design can be extended to more dilution levels and can be adapted to different dilution ratios . concentration errors in each stage propagate to the next level so that a careful design and fabrication of this structure are essential . while the residence time of drug molecules in the diluter branches must be long enough to assure complete mixing , the residence time in the incubation chamber must be as short as possible to avoid unwanted interference between neighbouring drug streams . the design requirements for the diluter and the incubation chamber are therefore strongly interrelated , and an optimization is necessary . fig3 shows a schematic of the incubation chamber . the five individual streams are flowing from the dilution stage into the chamber , where the cells 4 are immobilized and brought into contact with the drugs . the boundary conditions for the chamber design are as follows : 1 . mixing of the adjacent streams within the incubation chamber should be minimized . 2 . the flow rates q 1 to q 5 within the incubation chamber should be the same to provide an equal width of the drug streams in the incubation chamber . difflusion in the chamber causes a widening of the concentration profiles inside the incubation chamber . the maximal diffusion length , which is still acceptable , is given by the maximum time that is allowed without mixing to occur is given by for the device described in this disclosure , the minimum flow rate inside the incubation chamber should be at least 5 . 88 μl min - 1 taken all design parameters into account ( a = 200 μm ; d = 10 − 9 m 2 s - 1 for a typical biomolecule ; a = 0 . 35 mm 2 ; 1 = 1 . 4 mm ). fig4 ( a ) shows a schematic of the diluter : for both inputs , the buffer solution and the drug stock solution are directly connected to the ports 1 and 5 thus providing 0 % and 100 % streams . the diluter is realized as a cascading structure with three stages that mix the two solutions to the desired concentrations and connect these to the ports 2 to 4 . 2 . the output flow rates of the dilution stage q 1 to q 5 are equal ( normalized to 1 in this discussion ). 3 . the drug concentrations of q 1 to q 5 should be c5 = 10 c4 = 100 c3 = 1000 c2 ; c1 = 0 leading to relative concentrations of 100 %, 10 %, 1 %, 0 . 1 %, 0 % of the drug stock solution . the mixer structure has been designed using a lumped - element , equivalent - circuit model , in which each channel segment is represented by an electrical resistor . the individual flow rates and the resulting resistances of each branch can be determined by solving the linear system of equations derived from the equivalent circuit using kirchhoff &# 39 ; s theory . the flow rate corresponds to an electrical current and the flow resistance to an electrical resistance . the individual flow rates can be calculated using kirchhoff &# 39 ; s nodal rule as shown in fig4 ( b ). for the diluter output stream q 2 with a normalized flow rate of 1 , the ratio of the both incoming streams is 9 : 1 leading to a flow rate in the branches of 0 . 9 and 0 . 1 , respectively ( at each node the sum of the incoming currents equals the outcoming current ). the flow rates in the other branches can now be calculated bottom - up . the results are shown in fig4 ( b ). to achieve complete mixing in each branch of the diluter , a minimum residence time has to be assured . this condition is met for the overall system , if it is fulfilled for the mixing branch with the highest flow rate and the shortest channel length ( marked with the grey box in fig4 ( b )). if mixing can be guaranteed in this branch , the liquids in all other channels will be completely mixed as well . the flow rate of the mixing channel can be calculated by first determining the flow rates in the diluter output streams q 1 to q 5 . as all five branches of the diluter have the same flow rate and all drug streams are directed into the incubation chamber , the outlet flow rates q 1 to q 5 can be determined by with q chamber as the minimum flow rate in the chamber . the flow rate and the time required for a complete mixing can then be calculated by with q channel as the flow rate of the shortest channel , and the minimum required length of the channel to ensure complete mixing is then 2 . 7 mm for the given parameters ( a channel = 0 . 01 mm 2 ; q1 - 5 = 1 . 176 μl min1 ; q channel = 1 . 305 μl min - 1 ; t channel = 1 . 25 sec ). to increase the robustness of the system , the channel length was designed to be 6 mm . due to the required length , the channels are realized as meander - shape structures on the 2 × 2 cm 2 micro - fluidic chip . after the flow rate in each branch has been determined , the required resistance values can be analytically calculated using kirchhoff &# 39 ; s mesh and nodal rules . then , the electrical network can be translated back to a fluidic network , and the desired channel lengths can be determined . different flow resistances in the branches can be achieved by adapting the length of the channel segments ( flow resistance rl ˜ channel length l ). to assure reproducible mixing ratios even in the event of fabrication imprecisions , the cross - sections of all channels on the chip are identical . consequently , the only variable parameter is the channel length , however , the fabrication - induced variations are relatively small for this parameter . the silicon chip was fabricated in silicon - on - insulator technology ( 5 - μm device layer , 1000 nm silicon oxide , 380 - μm silicon handle wafer ) using combined front - and back - side etching ( fig5 ). first , five arrays of 200 orifices featuring 5 μm diameter were etched 5 - μm deep into the silicon from the front side by reactive ion etching . due to the required resolution , a chromium mask was used to photo - lithographically pattern a 1 . 8 μm thick photo resist layer ( s1818 , shipley , usa ) that serves as an etch mask . then , the back side of the wafer was patterned using 1000 nm pecvd silicon nitride as an etch mask for the wet - chemical etching . this etch mask has been structured by lithography and rie to define the membrane position . the 5 - μm thick silicon membrane underneath the orifice - array was formed by anisotropic etching 14 in 6 molar koh at 90 ° c . from the backside . the etching stops at the intermediate thermal silicon oxide , which was then removed using 10 % aqueous hf solution 15 to fully release the membrane and to open the orifices . the fabrication was completed by dicing the wafer into single chips . the diced chips were finally mounted on a flexible film ( face down ) and embedded in pdms by a casting procedure that will be described below . the micro - fluidic network was formed in a second chip which is fabricated in pdms by casting from a silicon mold featuring 100 - μm - high su - 8 structures . the fabrication process was as follows : after dehydration of the silicon wafer , the su - 8 ( su - 8 50 , microchem , usa ) was spun onto the wafer ( 1250 rpm ) and a two - level soft - bake ( 60 ° c . for 1 min , 95 ° c . for 75 min ) was performed on a hotplate to evaporate the solvents and to harden the photo resist . the hotplate was switched off after the bake to let the wafer cool down slowly . then , the uv - exposure in the mask aligner ( energy dose 600 mj / cm 2 ) was done to transfer the desired fluidic pattern from a typesetting film mask ( 8 μm resolution ) onto the wafer . the postexposure bake was carried out at 65 ° c . ( 1 min ) followed by 95 ° c . ( 45 min ), before the wafer was developed in microchem &# 39 ; s su - 8 developer for 10 min and washed with isopropanol . the fabrication was completed with the hard bake at 150 ° c . to achieve a better mechanical stability . the pdms replica mold was first pre - treated with the surfactant triton - x 100 ( 0 . 05 % in water ), which was applied by spin coating at 1000 rpm and then dried at ambient temperature . the surfactant is needed to facilitate the mold release of the pdms . then , the pdms ( sylgard 184 , dow corning , usa ) was prepared with a weight ratio of 10 : 1 for component a and b followed by degassing in a vacuum chamber for 30 min . the pdms was finally poured onto the wafer and cured at 60 ° c . for 4 hours . after removing the pdms layer from the master , the cast was rinsed thoroughly in warm water to remove triton - x residues that might prevent bonding and was cut into single chips . the silicon chip and the micro - fluidic pdms chip have dimensions of 7 × 5 mm 2 and 20 × 20 mm 2 , respectively . to prevent leaking of drugs through a cleft between these two devices , a tight seal between the silicon chip and the micro - fluidic cover is necessary . for that reason , the silicon chip was embedded into a pdms support to form a flat surface . the chip was first placed upside - down on a flexible polypropylene film , then , the cavity underneath the membrane was sealed by a 3 × 3 mm 2 teflon ( ptfe ) bolt , which was pressed against the chip . the pdms was poured around the chip and cured for 4 hours on a hotplate at 60 ° c . finally , the bolt was released and the plastic film was removed from the front side leaving the silicon chip seamlessly embedded in the pdms . to assemble the complete device , the pdms micro - fluidic unit was irreversibly bonded onto the embedded silicon chip after oxygen - plasma activation for 30 sec , 100 w . pipette tips ( 1 ml , roth ag , germany ) were used to fill the incubation chamber with the cell suspension . for the drug - screening experiments , a stepper - motor - driven syringe pump ( picoplus , harvard apparatus , usa ) was used to provide the required flow rates . two glass syringes ( ils gmbh , germany ) with volumes of 250 μl and 1000 μl to provide a flow - rate ratio of 1 : 4 of the drug stock and buffer solution were connected via dispensing needles ( 1 mm diameter , panacol , germany ) to the micro - fluidic device . before the cells could be loaded , the assembled overall device was cleaned with ethanol and exposed to an oxygen plasma at 80 w for 30 min to render the surface of the pdms less hydrophobic . directly after removing the device from the plasma furnace , the incubation chamber was coated with the adhesion - mediating protein laminin - 1 ( 20 μg / ml in tbs , sigma aldrich ) for improved cell adhesion . the chip was then incubated for 30 min , 37 ° c ., 5 % co2 before washing with tbs ( tris - buffered saline ). during the course of the experiments , a normal human dermal fibroblasts ( nhdf ) stock was cultured ( promocell , germany , c - 12300 ). before each cell loading , the medium was removed from the fibroblasts and the cells were washed with tbs . then , 0 . 25 % trypsin in medium ( invitrogen switzerland , 06354 ) in dmem ( invitrogen , 21885 - 025 ) was added ( 3 min , 37 ° c .) to detach the cells from the surface of the petri dish . the trypsin reaction was stopped with dmem containing 10 % fbs ( fetal bovine serum , sigma , f1051 ) ( at least 3 times the amount of trypsin ) and was then centrifuged at 1500 rpm before the supernatant was removed from the cells , and fresh medium was added . the cell clusters were then detached from each other by gently pipetting the cell suspension back and forth . the cell suspension was filled into a pipette tip , which was connected to one of the inlets of the cell loading ports . as the liquid level in this loading port was higher than in the other , empty one , a hydrostatic flow of cells into the incubation chamber was generated . the hydrostatic pressure difference between the inside and the outside of the incubation chamber also induced a minute flow through the orifices , so that single cells were trapped and were immobilized on the orifices . the cells were immobilized in five separate colonies of 200 cells each , so that the system provided a defined number of cells and a homogeneous cell density ( fig6 ). due to the larger specific density of the fibroblasts , the cells tend to sediment in the loading pipette . as a result , the cell concentration decreased permanently during the loading process until finally clear medium flowed through the chamber . as soon as all the orifices were occupied by cells , the remaining excess cells were , therefore , washed away . in fact , the cells were only retained in the chamber due to the pneumatic anchoring through the orifices . a control experiment using chamber without orifices yielded the result that no cells remained in the chamber . during cultivation , the loaded device was placed in a petri dish , which was filled with 2 ml of medium to prevent the drying out of the cells in the incubation chamber . the medium was exchanged once a day by hydrostatic flow using a medium filled pipette tip connected to the cell loading port . the performance of the drug diluter was first validated qualitatively using blue food color . for this experiment , the micro - fluidic device was bonded onto a glass microscope slide to be able to monitor the different color intensities under an inverted microscope . as calculated by our model , the flow rates were set to a ratio of 1 : 4 for the drug inlet and the buffer inlet at a total flow rate of 1 . 875 μl / min . fig7 ( a ) shows a micrograph of the diluter with the three mixing stages 18 . the mixing of the color and the buffer solution with a dilution ratio of 9 : 1 at each node could be qualitatively observed . after mixing , the drug and the buffer flowed through the long meander - shape channels 19 , which facilitated complete inter - diffusion . when entering the incubation chamber 8 , all drug streams were fully mixed , and five laminar streams of equal width through the chamber could be observed . at the entrance of the chamber , the streams were completely separated from each other ; further down a small degree of diffusion between the streams in the chamber could be observed as expected . however , the streams remained clearly separated and no major inter - diffusion between the neighbouring zones could be observed . moreover , the cell beds were spaced at a large enough distance and there was no concentration gradient over one of the cell beds ( fig3 ). for a quantitative evaluation , an aqueous 100 - μm fluorescein solution ( di - sodium fluorescein , sigma aldrich ) was filled into the drug inlet , and distilled water was filled into the buffer inlet . the fluorescence intensity was measured using a modified inverted epi - fluorescence microscope with a photo - multiplier module ( pmt h5784 , hamamatsu photonics , japan ) attached to the camera port the light emission from the chip was first spatially discriminated using a 1 - mm pinhole and filtered using a 525 - nm metallic interference filter ( edmund optics , usa ). fig7 ( b ) shows a plot of the calculated and the experimentally determined relative fluorophore concentrations . the graph shows that the fluorescence intensities produced at the outputs of the dilution cascade correspond very well to the desired concentrations . as the dilution of the different concentrations was achieved by a cascading structure , the deviation between the desired and achieved concentrations become larger from stage to stage yielding a maximum relative mismatch of 30 % for the 0 . 1 % dilution stage . however , this variation can be attributed to geometrical imprecisions in the micro - fluidic network as a consequence of the low resolution of the photolithographic mask . with a standard chromium mask , significantly better result is expected . nhdfs were chosen for the cell - adhesion and drug - screening experiments for several reasons : like most cells , fibroblasts only adhere to a surface if all culturing - conditions are met . but fibroblasts have the additional advantage that they change their shape to a triangular form upon adhesion , and after adhesion , fibroblasts start to divide when they are in a healthy state and are well supplied with all necessary nutrients . these characteristics allow for a convenient visual observation of the cell status . fig8 shows a micrograph of immobilized fibroblasts after 6 days in culture inside the 0 . 5 - μl incubation chamber . although the fibroblasts were immobilized on the orifices during the loading step and adhered to the laminin - coated surface , the cells expectedly began to migrate away from the orifices already after one day in culture and formed a homogenous cell layer . after 6 days in culture , a confluent layer of cells inside the incubation chamber was observed . this behaviour is desired because cell immobilization is only required during the loading phase to obtain a defined reproducible and homogenous cell population in the incubation chamber . once the initial population has been successfully established the cells should freely proliferate to form a confluent layer . to mimic a typical drug - screening procedure , the absorption of a fluorescent cell tracker by immobilized nhdfs from differently diluted streams of the fluorophore was studied . before the incubation , the chamber was coated with laminin - l ( 20 μl / ml in tbs ) for 30 min before cell loading . cell preparation and loading was performed as described in the experimental section . in this experiment , the dilute ion and cell exposure process was started already 30 min after immobilization . green cell tracker ( celltracker green cmda c2925 , molecular probes ) with a stock concentration of 100 μm was diluted to 10 μm , 1 μm , 0 . 1 μm and 0 μm with medium in the diluter stage . the cells were exposed to the five laminar streams of different concentrations at a total flow rate of 1 . 25 μl / min for 20 min . then , the syringe pump filled with the cell tracker solution was stopped , while the second pump with the medium continued operation to flush the chamber . the presence of the cell tracker was optically monitored as shown in the fluorescence image in fig9 ( a ). the concentrations increased from the left to the right . a correlation between the cell tracker concentration and the fluorescence intensity in the cell beds was observed . a more quantitative analysis is shown in fig9 ( b ) and was performed by image analysis of the acquired digital fluorescence images using the lspix - 5 . 1 ( national instruments of standards , usa ) software package . the average brightness of a rectangular area over each of the five cell beds comprising 64000 pixels was determined and plotted for each drug stream . while the higher - concentration streams produced significantly different fluorescence intensity in the cell beds , the 0 - μm and 0 . 1 - μm streams produced more fluorescence than expected . we attribute this to accidental contamination of the low - concentration streams with the cell tracker during starting the drug pump , which might have led to an intermittently increased drug concentration in streams 1 and 2 before a steady state was established . fig9 ( b ) also illustrates that the absorption of the dye in the cell caused a non - linear relationship between the cell tracker concentration in the stream and the corresponding cell fluorescence intensity ( note that the drug concentrations are logarithmic ). no major cross contamination between the neighbouring streams and cell beds was observable , so that the system met all requirements for a fully integrated cell - screening system . a combination of a micro - machined cell patterning and immobilization chip with online sample dilution over three orders of magnitude for cell - screening experiments was presented . by combining a small silicon chip for cell immobilization with an elastomeric micro - fluidics structure , a hybrid device featuring the advantages of precision silicon micro - machining and low - cost polymer replication techniques was fabricated . this device allows for arranging defined number of cells in a regular array , which improves the reliability of the experiment and allows for applying statistical methods . the integration of a micro - fluidic dilution cascade reduces both , the reagent consumption and the preparation time . a successful cell immobilization was achieved within 30 sec and cells were incubated in these devices for 6 days without observing reduced cell proliferation . the diluter stage was validated using a fluorescent dye , and a prototype screening experiment was performed using nhdfs and a fluorescent cell tracker . this shows that all the necessary procedures required for such an assay can be integrated in one system . 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