Patent Application
20250102505
The present invention relates to methods for a locally predetermined immobilization of one or more (e.g. single) cells under fluid flow conditions, in the context of a bioassay (biological assay), comprising for each (single) cell the step of trapping said cell in a reticular structure that is arranged in a microfluidic channel of a microfluidic device and extends in the directions transverse to a direction of fluid flow. Further, the present invention relates to respective bioassays comprising the step of (individually) immobilizing one or more (single) cells under fluid flow conditions according to said methods, e.g. assays for determining the kinetics of the binding of a ligand to the immobilized (single) cells. Furthermore, the present invention relates to respective microfluidic devices.
Major parameters of drugs and their performance are their likelihood to interact with the target (association rate) and the rate at which the complex will fall apart again (dissociation rate). This information about a molecular complex goes far beyond a general binding confirmation and leads towards pharmacokinetics in patients.
Kinetic parameters are classically determined in vitro with a purified target of the drug, e.g. by microfluidic biosensor instruments. However, it is known that the binding behavior of purified targets differs from targets in their natural setting embedded in cellular structures. The leading class of drugs in development for the treatment of all kinds of diseases, for instance of cancer, are antibodies. Many antibodies are developed to target cell surface markers and differentiate between healthy and abnormal cells.
To characterize the in situ binding kinetics and specificity of potential drugs in an early stage of development, it is desirable to immobilize cells in a microfluidic device. The measurement of interaction kinetics without technical artefacts entails the transport of the analyte (drug) to and from the target in a flow with a flow velocity that is preferably at least several 10 μm per second. Further enhanced analytic precision, however, may only be achieved with even significantly higher flow rates of even more than 1 mm per second.
A high flow rate around the cell ensures maximum supply of analyte and avoids the technical reduction of the measured association and dissociation rates by mass transport limitation. Furthermore, high flow rates are indispensable for the immediate removal of analytes as soon as an unbinding event is taking place, preventing rebinding which would result in apparently slower dissociation rates. The present concentration of e.g. antibodies on the cell surface can be determined by following the fluorescence of a label attached to the antibody via a microscopic setup above the immobilized target.
The immobilization of cells in a microfluidic device has no commercial platform to date. Experimental approaches applied include (i) specific capture via antibodies against cell surface markers; (ii) growing of adherent cells directly on a sensor; (iii) cell fixation; and (iv) mechanical barriers to trap the cells in flow.
Specific antibodies have the disadvantages of only being applicable to very specific cells and a rather low retention potential. It has turned out to be very challenging to make use of high flow velocities, when requiring a reliable and stable localization of the target cells. The same applies to adherent cells grown on the sensor surface, as such cells only resist rather slow fluid flow and are limited to adherent cell types. The fixing of adherent cells after growing them might improve their stability on the sensor. However, only kinetics with very limited flow rates have been published to date, indicating that no higher flow could be applied. Additionally, fixing of cells prevents any live-cell response measurements. Immobilization of cells by a mechanical trap can be applied to both adherent and suspension cell types, broadening the application. To date, polydimethyl-siloxane (PDMS) traps have been structured via soft lithography in microfluidic channels. These traps are large barriers and distort the flow before and after cell loading, reducing the fluid velocity that can be applied directly at the cell to about 50 μm/s. Cell loading and observation has only been published in slow flow rates up to a maximum of about 20 μl/min, which corresponded to fluid velocities of to-digit μm/s up to the low mm/s range, depending on the channel cross sectional area. Other technologies have been applied to cell sorting structures, which reduce the flow around each cell, as well as to cell cages for droplet deposition from the top, i.e., no flow is applied.
To determine kinetic rates of drug candidates as close as possible to the in vivo situation, it is necessary or desirable to measure their interaction with targets in situ, e.g. cell surface markers integrated in the native cellular membrane. From a technical point of view, high flow rates are required or desirable to measure real-time kinetics of molecular interactions in a microfluidic channel.
Accordingly, the technical problem underlying the present invention is the provision of means for reliable immobilization of cells, preferably single/individual cells, in a microfluidic environment under fluid flow conditions at flow velocities that are e.g. suitable for accurate kinetic measurements. The solution to the above technical problem is achieved by the embodiments characterized in the claims.
In particular, in a first aspect, the present invention relates to a method for the individual immobilization of one or more (single) cells under fluid flow conditions, comprising for each (single) cell to be immobilized the step of trapping said cell by means of a cell trapping element, which comprises a reticular structure that is arranged in a microfluidic channel of a microfluidic device and such that it extends in the directions substantially transverse to the direction of a main fluid flow in the microfluidic channel.
The method of the present invention is usually performed in the context of a bioassay, e.g. a bioassay as defined hereinafter. As used herein, the term “individual immobilization” relates to the fact that according to the present invention, preferably one single cell is immobilized or trapped in one reticular structure each, i.e. by a separate cell trapping element. More than one cell can be measured (and compared) in accordance with preferred implementations of the present invention by either capturing the cells in separate cell trapping elements and/or by measuring the cells in the same cell trapping element sequentially. In the latter case, after measurement of a trapped cell, said cell can be released in a controlled manner by reverting the direction of the fluid flow for a short time. And with subsequent forward flow, the next (single) cell can be trapped in the same cell trapping element.
Further, the term “under fluid flow conditions” as used herein is intended to indicate that the method of the present invention is performed, and each (single) cell is immobilized or trapped, hydrodynamically, while a fluid is flowing through the microfluidic channel, where the fluid flow is particularly applied at the location of the trapped cell. In other words, different from some conventional approaches, the present invention does not try to avoid or reduce an effective fluid flow at the location of the cell in order to prevent washing the cell away from its location of immobilization. Instead, the cell is directly exposed to the fluid flow by holding the cell trapped in the reticular structure of the respective cell trapping element, while the reticular structure allows largely unimpeded fluid flow at the location of the trapped cell. According to the present invention, this is achieved by the making use of the specific structure of the cell trapping element applied, wherein the reticular structure is composed of a plurality of struts that define a plurality of windows (openings), which allow fluid to flow through said windows.
Thus, the reticular structure defines a mesh with a mesh size adapted to retain a cell but to support fluid flow directly around the cell and through the windows (openings) in the mesh. In that respect, the windows may at least partially allow fluid flow therethrough even while a cell is trapped (and thereby immobilized) by means of the reticular structure. It turned out that such an arrangement is specifically effective in supporting high flow rates at the location of the immobilized cell without detaching the cell, which was a problem in known immobilization methods. The reticular structure thereby keeps any impact on the fluid flow very low and allows fast buffer exchange around the entire cell. Thereby, the present invention effectively combines good buffer exchange with accurate kinetic rate determination, where in conventional approaches a course tradeoff was required. Specifically, in most conventional approaches a choice was made between good buffer exchange which could easily wash away the cell as well, and very low flow rates and cells sitting in low flow areas which limited the accuracy of kinetic rate determination.
In preferred embodiments, the fluid velocity of the fluid flow at the location of the reticular structure and/or at the location of the trapped cell is set to a value of at least about 1 mm/s, preferably at least about 5 mm/s, more preferably at least about 10 mm/s, even more preferably at least about 50 mm/s, most preferably at least about 0.1 m/s or even at least about 0.5 m/s. Preferably, the fluid velocity at a location of the trapped cell and/or at the location of the reticular structure is specifically set to a value in the range of about 0.07 to about 0.74 m/s.
In a preferred embodiment, the microfluidic channel has cross section transvers to the direction of the (main) fluid flow specifically at the location of the at least one cell trapping element and/or at a center location of the trapped cell such that a volumetric flow rate of the fluid flow is at least about 20 μl/min, more specifically at least about 40 μl/min and preferably in the range from about 40 μl/min to about 3.5 ml/min, more preferably in the range from about 100 μl/min to about 2 ml/min, particularly while applying a fluid velocity according to one of the preferred values described herein. Specifically, the microfluidic channel preferably has a cross section transvers to the direction of the fluid flow in the range of about 0.2·104 μm2 to about 20·104 μm2, more specifically in the range of about 1·104 μm2 to about 5·104 μm2.
In this context, the term “volumetric flow rate” is defined as the volume of fluid which passes through the microfluidic channel in which the reticular structure(s) is/are arranged per unit time. Further, the term “fluid velocity” (or “flow velocity”) is defined as the momentary velocity of an element of fluid at a given location. As used herein, the terms “trapping/trapped” and “immobilizing/immobilized” are used synonymously and relate to the act of retaining one or more cells, preferably a single cell, at a given location in the microfluidic channel under fluid flow conditions by way of the reticular structures arranged in said channel.
In another aspect, the present invention also provides a microfluidic device for the individual immobilization of one or more (single) cells under fluid flow conditions of a bioassay. While the method for individual immobilization of one or more (single) cells as described herein may apply this microfluidic device according to any one of the preferred implementations disclosed herein, the structural and functional details disclosed in connection with the method are preferably also implemented in the microfluidic device. Thus, the microfluidic device comprises a microfluidic channel supporting continuous fluid flow along a fluid flow direction. It further comprises the cell trapping element having the reticular structure and being arranged in the microfluidic channel such that the reticular structure extends in the directions transverse to the fluid flow direction. As already explained in connection with the method for immobilization, the reticular structure is composed of a plurality of struts that define a plurality of windows allowing fluid to flow through said windows, such that (single) cells can be held trapped by the cell trapping element under fluid flow conditions.
In one specific aspect, the present invention relates to a bioassay comprising the step of individually immobilizing one or more (single) cells under fluid flow conditions according to the method of the first aspect of the present invention. Respective bioassays are not particularly limited and include any assay in which the immobilization of (single) cells under fluid flow conditions might be of interest. Such assays include e.g. assays for determining the kinetics of the binding of a ligand to the immobilized (single) cells. By way of example, such assays include studies of the binding kinetics of an antibody to a structure on the surface of the cells, or the binding kinetics of a drug or drug candidate to the cells. Such kinetic binding assays are known in the art. In specific embodiments, respective assays include the detection of a detectable marker, e.g. a fluorescent marker, that is bound to the cells.
In other embodiments, respective bioassays in which a physiological response of the immobilized cells is triggered, and said response, e.g. the release of a certain factor, is measured either directly on the cells or with the help of a biosensor in the immediate vicinity of the immobilized cells, or a biosensor lying downstream of the immobilized cells. The bioassay may assess and quantify cell viability, cell morphology, internalization of cell surface structure, release of vesicles and/or molecules of cellular origin, intercellular signaling, and/or changes in cellular metabolism and/or enzymatic activity in response to a stimulus presented to the cell in the fluid flow.
Preferably, a diameter of each strut of the reticular structure is in the range of about 0.2 μm to about 10 μm, preferably in the range of about 0.5 μm to about 5 μm, more preferably in the range of about 1 μm to about 3 μm, most preferably in the range of about 1.5 μm to about 2.5 μm. In this respect, a diameter of a strut is understood to define the average extension/diameter of a cross section of the respective strut perpendicular to its longitudinal extension. Specifically, in case of a circular cross section, the diameter of the strut is understood as the diameter of said cross section. Otherwise, e.g. in case of an irregularly shaped cross section, the diameter of the strut may be the diameter of an area equivalent circle. This structure combines a efficient combination of high cell trapping reliability and minimum distraction of fluid flow.
It has turned out to be specifically preferred if a size of each window along its respective maximum extension is not more than about 90%, preferably not more than about 80%, more preferably not more than about 70%, most preferably not more than about two thirds of a diameter of the trapped cell and/or wherein a size of each window along its respective minimum extension is at least 10%, preferably at least 20%, more preferably at least 30%, most preferably at least one third of the diameter of the trapped cell to be trapped. Thus, specifically with regards to the device, is it preferred that a size of each window along its respective maximum extension is at most about 15 μm, preferably at most about 12 μm, more preferably at most about 10 μm, most preferably not more than about 8 μm and/or wherein a size of each window along its respective minimum extension is at least about 1 μm, preferably at least about 2 μm, more preferably at least about 5 μm, most preferably at least about 7 μm.
It is also preferred that the struts of the reticular structure of the at least one cell trapping element form at least 4 windows, more preferably at least 6, even more preferably at least 9, most preferably at least 12 windows, and/or preferably not more than 50, further preferably not more than 25, most preferably not more than 20 windows.
Preferably the reticular structure of the cell trapping element defines a concave cell receiving surface facing an incoming fluid flow, wherein a smallest extension of the concave cell receiving surface transverse to the fluid flow is at least about 1.1 times, preferably at least about 1.2 times, more preferably at least about 1.3 times, even more preferably at least about 1.4 times, most preferably at least about 1.5 times a diameter of the trapped cell. This ensures a quite reliable cell trapping and enhances that chance that a trapped cell remains trapped during a bioassay. Alternatively or additionally a smallest extension of the concave cell receiving surface transverse to the fluid flow is at most about 10 times, preferably at most about 5 times, more preferably at most about 3 times, most preferably at most about 2 times the diameter of the trapped cell. This ensures minimal impact on fluid flow (around the cell) and high selectivity to trap individual cell.
In a preferred embodiment, a gold layer may be provided to cover part of an inner surface of a microfluidic channel wall, wherein the reticular structure of cell trapping element is attached to the gold layer. This gold layer may thereby improve the stability of the mechanical attachment of the reticular structure to the microfluidic channel wall. Additionally, the gold layer may serve as an electrode for bioassay making use of the efficient immobilization of one or more cells, preferably a single cell.
Thus, in one aspect the present invention also provides a bioassay comprising the steps of a method of individually immobilizing one or more (single) cells under fluid flow conditions according to any one of the aspects described herein; and determining the kinetics of the binding of a ligand to the immobilized (single) cell(s).
Preferably, the bioassay comprises an optical excitation and/or an optical measurement (spectroscopic analysis, such as fluorescence spectroscopy), wherein light is emitted to or from a location in or at the trapped cell along a measurement light path transverse to the direction of fluid flow, and wherein the reticular structure has the shape of a segment of the surface of a general cylinder with a cylinder axis parallel to the measurement light path and with a concave face of said segment facing the incoming fluid flow.
Thus, in a preferred implementation of the device, the reticular structure has the shape of a segment of the surface of a general cylinder (or prism) with a cylinder axis parallel to the measurement light path and with a concave face of said segment facing the incoming fluid flow. In this respect, the “general cylinder” means a geometric shape which is defined as a surface consisting of all the points on all the lines which are parallel to a given line (e.g. the cylinder axis) and which pass through a fixed plane curve in a plane not parallel to the given line. In one example which may be implemented as a preferred embodiment, that fixed plane curve may be circle and the resulting cylinder may specifically be a circular cylinder. Specifically, the cylinder axis may be perpendicular to the plane of said fixed plane curve. However, the “general cylinder” is not limited to a circular cylinder. Instead, the reticular structure may also follow the shape of a segment of an elliptical cylinder surface or a segment of the surface of a prism, which may result from a polygonal fixed plane curve. Regardless of the actual details of the shape of the cylindrical segment, it is preferred that said segment has one concave side (e.g. the side facing the inner side of the cylinder) that faces the incoming fluid flow. Thereby, the concave face of the cylindrical surface supports an improved capability of capturing and holding a cell while maintaining the flow condition in the fluid. Specifically, if the fluid flow a rather laminar (and not too turbulent) the cell may sit quite stable inside the concave surface. Specifically, with the cylindrical axis being parallel to the light path, the cell can be stably held at an optical observation position on the light path, while at the same time the reticular structure does not mask up the observation position. Instead, the excitation and/or measurement light can be sent along the light path without being blocked by part of the reticular structure.
In one preferred embodiment of the present invention, the reticular structure is attached with one end (section) thereof (e.g. the lower end) to an inner surface of a first microfluidic channel wall (e.g. a bottom/base wall of the microfluidic channel). This attachment is sufficiently stable to keep this first end section connected even during the desired fluid flow conditions. In this preferred embodiment, the cell trapping element further comprises at least one support beam having a first end attached to another end (section) of the reticular structure (e.g. an upper end). A second end of the at least one support beam is attached to the inner surface of the first microfluidic channel wall at a position downstream relative to the position where the reticular structure is attached. Thereby, the support beam supports the second (e.g. upper) end section of the reticular structure so as to stabilize the whole reticular structure even under the desired fluid flow conditions.
In an alternative implementation, the reticular structure is attached with one end (section) thereof (e.g. the lower end) to an inner surface of a first microfluidic channel wall, such as a bottom wall (base wall), and with another end (section) thereof (e.g. an upper end) to an inner surface of a second microfluidic channel wall, such as a top wall (cover) of the microfluidic channel substantially opposite to the first microfluidic channel wall. Thereby, each of opposite end sections of the reticular structure can be stable held at opposite walls of the microfluidic channel, thereby ensuring a high stability of the whole cell trapping element even under the desired fluid flow conditions. Preferably, the reticular structure of the cell trapping element is compressed between inner surfaces of the first and second microfluidic channel wall. Specifically, the reticular structure can be initially manufactured with a height that is larger than the height of the microfluidic channel (i.e. the distance between first and second microfluidic channel walls) before said opposing microfluidic channel walls are assembled, so that the reticular structure is clamped between these channel walls (e.g. the bottom and top surfaces of said channel. Specifically, the height of the reticular structure may initially be 10% to 30% larger than the height of the microfluidic channel.
Typical cell diameters for eukaryotic cells may range, dependent on the particular type of cell, in the region of 10 to 20 μm. However, some eukaryotic cells, e.g. yeasts or erythrocytes, can be smaller down to a few μm, and certain cell types are larger, e.g. human oocytes (typically 120 μm) or plant cells, which can extend over several 100 μm. Furthermore, prokaryotic cells can be between a few μm to as small as having a diameter below 1 μm. In this context, the type of cell to be immobilized in the methods of the present invention is not particularly limited in any way, including prokaryotic and eukaryotic cells of any kind.
Typical absolute values for a diameter of the half-cylindrical reticular structure may be in the range of 10 to 30 μm. However, depending on the type of cell to be immobilized, said diameter can range from 1 to 250 μm.
Preferably, the microfluidic channel has a substantially rectangular cross section transverse to the (main) fluid flow direction. In that respect, preferred values for a height of the microfluidic channel range from about 30 μm to about 75 μm. Moreover, preferred values for a width of the microfluidic channel range from about 500 μm to about 1000 μm, e.g. about 750 μm. Specifically, depending on the type of cells that should be immobilized, the height of the microfluidic channel may even be larger than 100 μm and may even extend over several 100 μm. Such relatively large channel heights may be specifically desirable if human oocytes of plant cells are to be trapped, for example. In some embodiments, it is preferred even for such (relatively) large channel heights that the reticular structure extends over the whole height of the microfluidic channel, so that the reticular structure can be supported by the channel walls opposing in height direction. Moreover, in cases with a substantially generally cylindrical shape of the reticular structure having the cylinder axis extending substantially in height direction, the extension of the reticular structure over the whole height can ensure that a trapped cell does not easily escape along the cylinder axis at one (open) end of the reticular structure, without requiring to provide a strut structure at such an open end as an additional barrier for the cell. Avoiding such a strut barrier specifically ensure to avoid masking excitation and/or measurement light that can be sent substantially parallel to the height direction and/or the cylinder axis.
A large channel height, optionally even combined with an analogously large extension of the reticular structure in height direction, can even be advantages in cases where (relatively) small cells are to be immobilized. Even if the reticular structure is adapted in width direction (i.e. transvers to the direction fluid flow and perpendicular to the height direction), e.g. the direction of the smallest extension transverse to the fluid flow, in accordance with the diameter of the cells to be immobilized as described above, the reticular structure may extend in height direction over the whole height of the microfluidic channel. Thus, in some preferred embodiments, the reticular structure may be multiple times higher than wide. In these embodiments, the limited extension of the reticular structure in width direction can ensure a high selectivity in trapping (single) cells of a specific size, while the relatively large extension of the reticular structure in height direction may allow to hold multiple identical cells, preferably aligned along an axis that may be used as an optical axis for excitation and/or measurement light. This can be specifically advantages with regards to a high response sensitivity for bioassays in which a physiological response of the immobilized cells is triggered, and said response, e.g. the release of a certain factor, is measured with the help of a biosensor in the immediate vicinity of the immobilized cells, or a biosensor lying downstream of the immobilized cells.
In some embodiments it is preferred that the cell trapping element comprises a support structure such as the support beam described herein, even if the reticular structure extends over the whole height of the microfluidic channel and even if the reticular structure is supported by the walls of the microfluidic channel opposing in height direction. This is specifically advantageous for the mechanical stability of the reticular structures even in high flow velocities in cases of a high aspect ratios between the height and the width of the reticular structure, e.g. if the reticular structure is at least 5 times or even at least 10 times higher than wide.
In specific embodiments, the diameter of each strut of the reticular structure is 1.5 to 2.5 μm, preferably 1.6 to 2 μm, more preferably 1.7 to 1.9 μm, e.g. about 1.8 μm.
The reticular structures used in the methods of the present invention comprising rectangular windows (or meshes). Preferably, the length of the larger side of said rectangular windows is at most two thirds of the diameter of the cells to be trapped, and the length of the smaller side of said rectangular windows is at least one third of the diameter of the cells to be trapped.
In absolute values, the length of the larger side of said rectangular windows is preferably between 5 and 15 μm, and the length of the smaller side of said rectangular windows is preferably between 2 μm and 7 μm. However, depending on the size of the cells to be trapped, the larger side of said rectangular windows can be between 0.6 to 100 μm, and the length of the smaller side of said rectangular windows is preferably between 0.3 μm and 50 μm.
The reticular structures to be used in the methods of the present invention can be fabricated by photon-stimulated polymerization, such as 2-photon polymerization (2PP) as known in the art. Accordingly, the material said structures are made of include suitable photopolymers, such as negative tone IP-Dip or IP-Visio photoresigns, SU8, positive tone photopolymers and hydrogels.
The bottom surface of the microfluidic channel used in the methods of the present invention, i.e., the surface on which the reticular structures are adhered, are not particularly limited and include any surfaces on which respective structures can be fabricated by 2PP. Suitable materials in this respect include glass, silicon, metals (e.g. gold and/or chrome), ceramics and/or polymers such as polydimethylsiloxane (PDMS), polypropylene (PP) and/or SU8. In preferred embodiments, the bottom surface of the microfluidic channel, and/or the top surface of said channel, are made of an optically transparent material, e.g. glass, in order to allow for microscopic, e.g. fluorescence microscopic, observation and/or analysis of the immobilized cells.
Means for actually immobilizing cells using the methods of the present invention are not particularly limited and are known in the art. They include any means of loading the respective microfluidic channel/microfluidic device with a fluid containing cells. Suitable fluids include any suitable buffers and media which allow the suspension of cells in the microfluidic flow. These can be of natural (e.g. serum, blood, amniotic fluid) or artificial source or mixes thereof. Artificial fluids include buffered salt solutions such as phosphate buffered saline (PBS) or culture media which keeps the pH of the cell suspension stable while supplying required nutrients, e.g. the basic medium MEM or the more complex medium RPMI-1640.
The cell density of a respective fluid used to load the microfluidic channel/microfluidic device with cells can be in the range of 0.1-106 cells/ml. Lower cell concentrations are possible as well with respectively longer loading times or additional guiding structures towards the (single) cell traps to enhance the probability of the cell entering the trap. In a preferred embodiment, the guiding structure(s) may comprise a mesh wall extending substantially from bottom to top of the microfluidic channel such that it allows substantially unrestricted buffer flow while it guides any incoming cell towards the reticular structure of the cell trapping element. Thereby, cells which would have passed by the trap in a channel that is broader than the reticular structure of the cell trapping element are forced to enter the trap first and only surplus cells will pass the filled trap. This can ensure a considerable improvement of the cell trapping efficiency. In order to ensure passage of surplus cells, it is preferred that a gap is provided between the guiding structure and the reticular structure of the cell trapping element, which is larger than the diameter of the cells to be trapped.
The microfluidic devices of the present invention can comprise one or more of the cell trapping elements, each including a reticular structure. In cases where more than one cell trapping elements is present, they are preferably arranged so that the distances therebetween are at least not smaller than the size of the cell trapping elements (i.e. their extension in the direction of the distance therebetween). This ensure a limited distortion of the fluid flow and avoids blockade of cell flow at undesired positions between the cell trapping elements. Furthermore, in case of more than one cell trapping elements being arranged in multiple rows with respect to the direction of fluid flow, the respective structures in different rows are preferably arranged in an offset manner with respect to the different rows.
In preferred embodiments, the methods of the present invention are performed in vitro, i.e., they are not performed on the human or animal body.
As used herein, the term “about” is intended to be a modifier of the specified value of ±10%, preferably ±9%, ±8%, ±7%, ±6%, ±5%, ±4%, ±3%, ±2%, or ±1%. By way of example, the term “about 10 μm” is intended to encompass the ranges of 9 to 11 μm, preferably 9.1 to 10.9 μm, 9.2 to 10.8 μm, 9.3 to 10.7 μm, 9.4 to 10.6 μm, 9.5 to 10.5 μm, 9.6 to 10.4 μm, 9.7 to 10.3 μm, 9.8 to 10.2 μm, or 9.9 to 10.1 μm.
In the present invention, as a first requirement for immobilizing both adherent and suspension cells non-specifically, a mechanical trap system was focused on. First trap designs by photolithography included solid pillar structures. These pillar arrangements enable to stop cells in a microfluidic flow. However, they are quickly filled up by many cells which results in a total block of buffer flow, preventing any controlled kinetic experiment on top of the captured cells. Thus, a requirement for the trap being made up of a mesh structure was identified, and the reticular structures of the present invention realized as an ideal mechanical cell trap system.
In particular, the present in invention provides robust (single) cell traps fabricated in a microfluidic channel preferably by two-photon polymerization (2PP) that withstand a very high average flow velocity of 1 m/s or more. The traps may be manufacture by a direct laser written thin reticular structure allowing a good flow around the captured cells as well as optical measurements. Therefore, they enable the characterization of binding kinetics between active substances and cells. Preferably, the traps may have the shape of half a circular cylinder and capture the cells hydrodynamically.
Further details and advantages for preferred embodiments of the present invention will be described with reference the attached figures, which show:
In the examples of
While the left most cell trapping element in
The shape of the reticular structures with a substantially circular cylindric shape corresponding to a half-circular base was designed such as to well accommodate a spherical cell. A mesh-like structure was chosen as trap wall to allow unrestricted buffer passage. The windows 18 in the trap wall allow flow to pass through the traps and do not divert flow sideways. The window size was optimized to stop cells of >10 μm diameter in flow but allow smaller cell fragments to pass through and not clutter the trap. A good window size for ˜ 15 μm large Jurkat cells is 6×8 μm.
Structures of this size and complexity cannot be produced with standard photolithography methods. Therefore, the 2PP method was chosen for direct laser writing of the cell traps in an IP-Dip polymer with a Photonic Professional GT (Nanoscribe GmbH) (
Initial tests showed possible flow-dependent instability of the traps. Therefore, to increase the stability, the height of traps was designed to be slightly more than the channel height, which is leading to clamping of the trap between top and bottom cover slip of the biosensor chip as will be illustrated in more detail in connection with
In order to test the influence of support structures on the stability of the cell traps according to the present invention at different flow rates, cell traps having no, one, or two stabilizer bars extending from the top backside of the reticular structure to the surface of the bottom surface of the microfluidic channel in the direction of fluid flow, forming an angle with said bottom surface of about 45° were generated.
As shown in
| Number | Date | Country | Kind |
|---|---|---|---|
| 20176518.7 | May 2020 | EP | regional |
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/EP2021/064126 | 5/26/2021 | WO |
