Claims:WE CLAIM: 1. A micro channel network housing with rounded micro masts in Bio MEMS for handling bioparticles within a fluid flow in continuous manner, the device comprising: - at least two different inlets operative to receive fluid carrying bioparticles and buffer as input; - at least one inlet which is centrally located for sample input; - at least one outlet each for CTC (Circulating Tumor Cells)and WBC (White blood cells), separation disposed adjacent to each other; - at least one control outlet for administering the fluidic pressure in uniform manner; - a means for separation of bioparticles using the flow of fluid; and - a network inclusive of micro masts arrangement with inclined angles of the channels operational to appropriately direct the movement of said fluid, the micro channel being in a vertical and horizontal arrangement. 2. The micro channel network housing circular micro masts as claimed in the claim 1, where the said mechanism is configured utilizing Deterministic Lateral Displacement (DLD). 3. The micro channel network housing circular micro masts as claimed in the claim 1, where the aforesaid inlet operates along with other two buffer inlet streams as per the shear fluid flow rate of micro channel and restrains as well as redirects the fluid flow of stream at the neutral stages of channel containing bioparticles in said micro channel with side plane walls having definite width. 4. The micro channel network housing circular micro masts as claimed in the claim 1, where the aforesaid Micro channel Network directs the streamlines carrying target bioparticles along the Separation channel (horizontally). 5. The micro channel network housing circular micro masts as claimed in the claim 1, where the aforesaid Micro channel Network consists of distinctive inlets and outlets from which sample is injected and bioparticles are separated in accordance to the sample inlet from which a sample containing bioparticles for separation are frequently impregnated. 6. The micro channel network housing circular micro masts as claimed in the claim 3, where the two inlets i.e. from Buffer inlet-1 and buffer inlet-2 a sample without cells/ bioparticles is impregnated to minimize the fluid flow fluctuations due to which the cells/bioparticles mobility can be examined in a continuous manner for the separation of particles in the micro channel. 7. The micro channel network housing circular micro masts as claimed in the claim 1, where the aforesaid Micro channel Network is designed for various sized bioparticles allowed to traverse through specifically designed pattern of micro masts inside the micro channel for sample constituent separation by assimilating the geometrical parameters in network where separation occurs at each intersection terminals within the micro channel. 8. The micro channel network housing circular micro masts as claimed in the claim 1, where the Micro channel Network encompasses circular micro posts sized 12µm in diameter and gap of about 22 µm in between micro posts maintaining a tilted angle of 13.380 for the DLD array arrangement. 9. The micro channel network housing with rounded micro masts as claimed in the claim 1, where the fluid flow injected through the sample inlet as well as buffer inlets have laminar flow with Reynolds number less than or equal to 1. 10. The micro channel network housing with rounded micro masts as claimed in the claim 1, where the micro channel thickness i.e. the cylindrical height of micro posts is around 31µm. 11. The micro channel network housing with rounded micro masts as claimed in the claim 1, where the Micro channel Network is designed for separation of CTC from sample of Blood and other related infectious sample. 12. A method for separation of bioparticles in a fluid flow constitutes: a) providing in a micro channel network a patterned array of circular micro masts, orienting the streamlines as per the principle of DLD forcing the target bioparticle (CTCs) to percolate through a definite stream between the designed spacing between micro pillars leading to lateral shifting of target cells focusing on distinct isolation through designated outlets. b) precipitating a flow of fluid through aforesaid network; c) accomplishing bioparticles in said fluid flow directed along at most one direction as they outflow over said micro channel; d) introducing three inlets i.e.(Sample inlet, Buffer inlet-1 and Buffer inlet-2), a fluid sample of blood containing target bioparticles (CTCs) is insinuated to minimize the flow fluctuations through which the motion of bioparticles can be examined over and over for separation in said micro channel; e) aggregating the separated bioparticles as they leave through said outlet. 13. The method for separation of bioparticles from a fluid flow as claimed in claim 12, where the aforesaid network is constructed utilizing Deterministic Lateral Displacement (DLD). 14. The method for separation of bioparticles from a fluid flow as claimed in claim 12, where the aforesaid micro channel network directs the streamlines carrying target bioparticles along the Separation channel (horizontally). 15. The method for separation of bioparticles from a fluid flow as claimed in claim 12, where the aforesaid micro channel network encompasses circular micro posts sized 12 µm in diameter and gap of about 22 µm in between micro posts maintaining a tilted angle of 13.380 for the DLD array arrangement. 16. The method for separation of bioparticles from a fluid flow as claimed in claim 12, where fluid flow injected through the sample inlet as well as buffer inlets have laminar flow with Reynolds number less than or equal to 10. 17. The method for separation of bioparticles from a fluid flow as claimed in claim 12, where the micro channel thickness i.e. the cylindrical height of micro posts is around 31µm. , Description: MICROFLUIDIC CHANNEL HOUSING WITH ROUNDED MICRO MASTS FOR ENTRAPPING OF CIRCULATING TUMOR CELLS EMPLOYING DETERMINISTIC LATERAL DISPLACEMENT PRINCIPLE FIELD OF INVENTION The present invention caters to the domain of Biological MEMS and is based on the principle of Deterministic Lateral Displacement employed in a microfluidic channel housing rounded micro masts for isolation of Circulating Tumor Cells. More particularly, the disclosure specifically relates to the separation of cells/bio-particles in blood samples using a passive technique called Deterministic Lateral Displacement in a microfluidic channel which isolates the target cells, in this case Circulating Tumor Cells from White blood cells based on their size and mass under the action of hydrodynamic forces. BACKGROUND ART An unconventional growth in research has been identified in the topic of MEMS (microelectromechanical system) which emerges from silicon based integrated circuit technology that have remarkably brought convincing developments demonstrating innumerable breakthroughs technically, particularly focused in the field of biological/biomedical technology. Over the years, an enormous improvement in technical advancements have been observed in evolving areas of medical diagnosis from which separation/sorting of bio-particles has been receiving utmost consideration as the focus of research is relocating towards betterment in therapeutic approaches for interception of life threatening diseases. Various orthodox techniques of analysis in biological field essentially depend on the different techniques along with the level of experience of operators with actual traditionally known instruments for large scale operations that offer limited accuracy in adulterated samples for detection and also demand the utilization of samples at a superior rate. In the recent times, modelling and analyzing a microfluidic structure in biomedical field has been suitably tagged as an emerging research area that advances to develop with each passing year alongside the replication of modern evolution in analyzing a sample. Approaches in miniaturization of mechanism just as analysis tools at micro level generally termed as ‘Lab-on-a-chip’ have diminished complications at macro-scale level by employing definite miniaturized systems like micro valves, micro mixers, micro pumps, and micro channels. The analogous microfluidic devices for biological sample analysis necessitates less consumption of reagent, sample utilization at low level along with less cost whenever correlated with considerable conventional methods being utilized [1]. In the continually emerging field of Biological MEMS, sorting/separation of bio-particles from the sample has attracted particular interest in therapeutic and diagnostics applications like purifying of targeted particles/cells and regulating the bioassays of living particles or tissues for further study. For achieving the separation of bio-particles, a few primary characteristics of particles considered are geometrical shape, size, ability to deform, density, compressibility and many other related physical characteristics. Figure 1 shows the various types of processes available as of now which are frequently utilized for separation of bio-cells. Of all the different types of methods for separation, one of the methods is passive type where Pinched Flow Fractionation (PFF) is a technique for isolating particles/cells that involves slender micro channel termed as a “pinched segment” through which the particles/cells percolates in-line with the micro channel side wall and separates over various outputs on the basis of critical diameter. One of the competent methods for the separation of particles is Deterministic Lateral Displacement (DLD) where the cells/particles converge over an array of planned placements of posts. One type of the significant and structural physical principles in the field of microfluidics studies is hydrodynamic effect which includes inertial and dean rotation force that is operated based on the theory of fluid dynamics in micro-channeling network and utilized distinctly to establish effect of hydrodynamic force based separation. Other than the passive type there exits the active type, where one of the methods is Acoustophoresis that utilizes continuing wave in the form of ultrasonic waves that controls the movement of cells just like other techniques namely optical, dielectrophoresis, magnetophoresis and other active methods. An integral principle of acoustic separation/sorting employs ultrasonic waves to generate pressure gradients for separation/sorting. Magnetism principle can be used for the separation of bioparticles utilizing a technique named magnetophoresis where bio-cells are attached with the magnetic beads and separation occurs depending on the deflection of particles/cells due to magnetic susceptibility, fluid flow rate and size. Another technique is DEP (Dielectrophoresis) method, where the polarizable bioparticles/cells whenever suspended in the electric field of non-uniform pattern experience mobility produced due to Dielectrophoresis force directed by the actual electric field along with bioparticles electrical properties/solutions. In the present scenario of medical advancements where the number of patients diagnosed with cancer is on the rise, substantially researchers have been focusing on the trapping of CTCs from blood sample/blood stream as a fluid sample excision for diagnostic study of cancer at early stage. Naturally, CTCs are genuinely deficient, amounting to around 109 of cells in specific unit (ml) of bloodstream fluid of patients suffering from metastatic cancer. The definite action of Circulating Tumor Cells (CTCs) in the process flow of metastasis stands imprecise. Thus, to accelerate the diagnosis of cancer, study and treatment of tumor metastasis, characterization, trustworthy and effective techniques for enumeration of CTCs and sorting/isolation are being conducted at elevated levels. Diameter of Circulating Tumor Cells (CTCs) mostly lie in between the range of 12–20 µm, on the other hand in literatures it has been marked that the cell size range lies around 27 µm diameter approximately. For the separation of Circulating Tumor Cells (CTCs), utmost concerning and often utilized techniques were on the basis of label-less isolation of CTCs that includes deterministic lateral displacement (DLD), inertial focusing, microfluidic filters, dielectrophoresis (DEP), acoustics and Pinched Flow Fractionation (PFF). The above-mentioned techniques are operated on the basis of the fact that Circulating Tumor Cells (CTCs) are bigger as well as more massive than the regular sample of blood cells. On the other hand, the techniques have their own limitations based on their working. Though DEP (Dielectrophoresis) and Acoustics are the techniques which necessarily demand additional force fields along with longer process time although the other techniques like inertial focusing and microfluidic filters have shortcomings with respect to the issues of clogging which is one of the crucial part for the separation of bioparticles. Deterministic Lateral Displacement (DLD) is an extrusive technique within the utilized methods for particles sorting/separation in the modern era. Essentially it is size-based sorting/separation method which accomplishes the bifurcation of laminar fluid flow through circular masts that are allocated in an oblique array in the interior of the micro channel. An oblique array of masts leads towards competent displacement by the virtue of hydrodynamic forces that acts under the laminar flow due to the masts pattern. One of the important parameters of DLD technique known as critical diameter (Dc) helps to estimate the displacement of the suspended bioparticles in a micro channel through the oblique arrays of micro posts based on the angular changes done. As per the fluidic principle, if the diameter of bioparticle (Dp) suspended in micro channel is lesser than the critical diameter then there won’t be any streamline relocation in lateral mode and will lead towards an uneven trajectory of fluid flow. As per Figure 2, contrarily, if bioparticle’s width transcends critical diameter, the pattern of bioparticle fluid flow will resemble a substantially bumped fluid flow style. In this device as it utilizes Deterministic Lateral Displacement method, one of the geometrical parameters i.e. row shift fraction (e) representing the ratio of deviated displacement in between pillars (micro posts) and displacement from intermediate points of neighboring pillars (micro posts) is used. Considering the enactment of parallel row deviation of pillars (micro posts) in DLD device arrangement, a characteristic parameter termed as tilted (slanted) angle (?) is employed by systematized manifesting of the gradient of pillar (micro post). A period number (N) is also examined, that reveals about the pillars (micro posts) of row N + 1 which are in the equivalent location with each other in parallel manner as depicted in first row. The parameters mentioned above are related by Equation-1 that associates the parameters e with N and ?, e= 1/N=tan??(?)? (1) Critical diameter (Dc) of a Deterministic Lateral Displacement network could be approximated for Reynolds number (Re) at lower rate with the aid of Equation-2 that was determinative for the operation implicating Re < 1. D_c=1.4.g.e^0.48 (2) In the event when moderate throughput in between 1 2000 defines turbulent with no linear response amidst pressure drop along with fluid flow rate. However, in a flowing streamline channel having a laminar flow alongside Re beneath 1 at continuous micro-scale height grants effect of inertia at minimum value. During the experimentation in COMSOL Multiphysics a definite environment was created consisting independent inlets, planed side walls prognosis and multiple outlets for pressure alongside velocity profile of geometrical structure was designated with respect to the micro-channel structure with explicit velocity standards and mutual separation scale of cells/bio-particles (biological species like CTC and WBC) along with the properties of different parameters. Result Here, the responses for simulated micro channel model was determined after computational analysis of channel network in COMSOL Multiphysics 5.4 software which are conferred by representing a plot with alteration of distinctive fluidic sample parameters with variance in fluid flow rates. Particular crucial cell specification contemplated for bioparticle tracing with mass in the micro channel network, simulation parameter for cells shown in Table 3 which were referred from the articles/literatures related to Circulating Tumor Cells (CTCs). Table 3: Particular cell parameters contemplated for bioparticle tracing in micro channel network Bioparticle/Cell Class Diameter (µm) Mass (pg) White Blood Cell (WBC) 12 19.3 A cell extracted from an Ovarian Cancer Patient namely Circulating Tumor Cell (CTC) 27 36.8 SIMULATED RESULTS Simulation results of the determined micro channel structure where the conducting medium was taken as water contemplating at variance of velocity fluid flow rate Practicing one of the features of COMSOL Multiphysics 5.4 i.e. particle tracing feature, micro channel structure was examined by interjecting a sample of blood which contains Circulating Tumor Cells (CTCs) and WBCs of semi diameter 13.5 µm and 6 µm respectively through sample inlet with conducting fluid as water. Figure 6 demonstrates the response after simulation of the contemplated structure which evidently depicts a decisive sorting/separation of Circulating Tumor Cells (CTCs) at 4.75 m/s of sample velocity fluid flow rate. The velocity fluid flow rates of inlet-2 for buffer and inlet-1 for buffer were retained at 6.7 m/s and 4 m/s respectively. Simulation results of the determined micro channel structure where the conducting medium was taken as blood contemplating at variance of velocity fluid flow rate Practicing one of the features of COMSOL Multiphysics 5.4 i.e. particle tracing feature, micro channel structure was examined by interjecting a sample of blood which contains Circulating Tumor Cells (CTCs) and WBCs of semi diameter 13.5 µm and 6 µm respectively through sample inlet with conducting fluid as blood. Figure 7 demonstrates the simulation response of the contemplated structure which evidently depicts a decisive sorting/separation of Circulating Tumor Cells (CTCs) at 12.3 m/s of sample velocity fluid flow rate. The velocity fluid flow rates of inlet-2 for buffer and inlet-1 for buffer were retained at 6.7 m/s and 4 m/s respectively. Simulation results of the determined micro channel structure where the conducting medium was taken as blood contemplating mass fluid flow rate Practicing one of the features of COMSOL Multiphysics 5.4 i.e. particle tracing feature, micro channel structure was examined by interjecting a sample of blood which contains Circulating Tumor Cells (CTCs) and WBCs of semi diameter 13.5 µm and 6 µm respectively through sample inlet with conducting fluid as blood. The thickness of micro channel for all inlets was abbreviated to 0.01 mm to achieve Re fluid flow characteristics at higher rate. Figure 8 demonstrates the simulation response of the contemplated structure which evidently depicts a decisive sorting/separation of Circulating Tumor Cells (CTCs) at 4.3 x 10-6 kg/s of comparable superior sample mass fluid flow rate. Mass fluid flow rates of inlet-1 for buffer and inlet-2 for buffer were retained at 30.8 x 10-7 kg/s and 5.96 x 10-6 kg/s respectively. The structural model presented in a recent literature referenced in [1] where isolation of CTCs was obtained with appreciable purity at 150 µl/min that corresponds to 2.65 x 10-6 kg/s taking into account the density of blood sample as 1060 kg/m3. Therefore, response of the simulation result for the proposed micro channel in clear sense justifies the competent activity of the structural micro channel at high mass fluid flow rate as well as Reynolds number (Re number as 8.9). Comparative Analysis with respect to the Referenced Prior Art Steady performance of a microfluidic channel for distinct isolation of target particles/cells under high throughput operation is currently one of the prioritized objectives in this domain. However, in prior art i.e. in US specification US 9433880, particle separation was distinctly reported for two different nominal flow velocities i.e. 1.75 cm/s and 5.25 cm/s. It can be observed that the spiral design of the micro channel limits the nominal inlet velocity flow rate to a lower value making it unsuitable for high throughput operations. Moreover, the pressure distribution profile, as depicted in the prior art (US 9433880) appears to be non-uniform and non-linear throughout the spiral microfluidic channel profile. This presents a theoretical possibility of clogging near the outlets as the particles traverse in different pressure profile zones. Also in prior art, US specification US10324011B2, DLD arrays using triangular posts were used to concentrate CTCs at high flow rate of 10 ml/min. Use of triangular posts can lead to rupture of isolated CTCs. In an attempt to address the cited issues above regarding the prior art (US 9433880 and US10324011B2), the designed micro channel network introduces two specific improvements: The designed micro channel network achieves isolation of CTCs at a much higher inlet flow velocities of 4.75 m/s and 12.3 m/s, and high mass flow rate of 4.3 x 10-6 kg/s using simplified circular posts in DLD array thereby making the design better suited for high throughput operations and avoiding rupture of isolated CTCs. As per the COMSOL Multiphysics simulation result provided in Figure 9, the pressure variation inside the micro channel network is linear and an absolute uniformity of pressure profile is maintained near the outlets making the design immune to clogging issues. In another prior art, US specification US2015-0362413, separation of RBCs were reported at inlet fluid velocities of 200 µm/s and 1000 µm/s with various pillar shapes. On the contrary, the designed micro channel network displays capability of isolating a more critical diagnosis biomarker, CTCs at higher inlet fluid velocities of 4.75 m/s and 12.3 m/s promising higher throughput. Additionally, simplicity of rounded micro masts in the DLD array contributes towards ease of fabrication leading to less possibility of instrumental errors in designing DLD pillars. In yet another prior art, US specificationUS20150238963A1, separation was possible when the smaller particles range (size) between 5 µm -15µm and larger particles range (size) between 25 µm - 40µm for a sample flow rate of 2.5 mL/min – 3.0 mL/min. The designed micro channel network is capable of isolating two cells (CTCs and WBCs) having very small size difference (i.e. 12 µm and 27 µm) at equivalent sample mass flow rate of 4.3 x 10-6 kg/s. Plots and Tables: Plot of mutations inessential fluid flow parameters corresponding to the variation in sample velocity fluid flow rate Figures 10 – 13illustrates the mutation in fluid flow parameters particularly Reynolds number (Re), pressure (p), velocity magnitude and shear rate at contradistinctive outlets of contemplated micro channel device with variations in velocity fluid flow rate utilizing conducting fluid as water. Circulating Tumor Cells (CTCs)attains the separation adequately at velocity fluid flow rate for sample attributed at 4.75 m/s and further the linearity of the fluid flow characteristics is established as we alter the fluid flow rate facilitating from the sorting/isolation point. Furthermore, Figures 14 – 17 interpret the alteration of fluid flow parameters particularly Reynolds number (Re), pressure (p), velocity magnitude and shear rate at contradistinctive outlets of contemplated micro channel device with variations in velocity fluid flow rate utilizing conducting fluid as blood. Circulating Tumor Cells (CTCs) attains the separation adequately at velocity fluid flow rate for sample attributed at 12.3 m/s and the linearity of the fluid flow characteristics is established as we alter the fluid flow rate facilitating from the sorting/isolation point. Figure 10 depicts the mutations in Pressure (Pa), Figure 11 depicts the mutations in Reynolds number (Re), Figure 12 depicts the mutations in Shear Rate (1/s) and Figure 13 depicts the mutations in Velocity magnitude (m/s) on the three contrasting outlets corresponding to the alterations in fluid flow rate (m/s) of sample. The graph evidently displays the consistency in pressure (Pa) dissemination, Reynolds number (Re), Shear Rate (1/s) and Velocity magnitude (m/s) at contrasting outlets post the point of separation of Circulating Tumor Cells (CTCs) which was at 4.5 m/s fluid flow rate whenever conducting fluid was used as water Table 4: Measured Data regarding mutations in pressure (Pa) at contrasting outlets corresponding to the alteration in fluid flow rate (m/s) of sample with conducting fluid considered as water Sample Inlet flow rate Apex Buffer Inlet flow rate Bottom Buffer Inlet flow rate Apex CTC outlet Pressure (Pa) Central WBC outlet Pressure (Pa) Bottom Outlet Pressure (Pa) 4.75 m/s 4 m/s 6.7 m/s 429.39 104.05 -20.466 5 m/s 4 m/s 6.7 m/s 436.71 114.09 -25.549 5.75 m/s 4 m/s 6.7 m/s 478.22 136.96 -53.514 6 m/s 4 m/s 6.7 m/s 553.79 139.21 -75.15 6.75 m/s 4 m/s 6.7 m/s 610.97 148.56 -93.4 7 m/s 4 m/s 6.7 m/s 676.76 154.18 -91.401 7.75 m/s 4 m/s 6.7 m/s 851.26 165.92 -101.6 10 m/s 4 m/s 6.7 m/s 942.93 199.7 -88.554 12 m/s 4 m/s 6.7 m/s 1082.9 207.75 -132.54 Table 5: Measured Data regarding mutations in Reynolds Number (Re) at contrasting outlets corresponding to the alteration in sample fluid velocity flow rate (m/s) with conducting fluid considered was water Sample Inlet flow rate Apex Buffer Inlet flow rate Bottom Buffer Inlet flow rate Apex CTC outlet Cell Reynolds number Central WBC outlet Cell Reynolds number Bottom Outlet Cell Reynolds number 4.75 m/s 4 m/s 6.7 m/s 11.353 3.0624 10.815 5 m/s 4 m/s 6.7 m/s 11.785 3.183 11.205 5.75 m/s 4 m/s 6.7 m/s 12.149 3.2286 11.518 6 m/s 4 m/s 6.7 m/s 12.637 3.3197 11.559 6.75 m/s 4 m/s 6.7 m/s 13.159 3.4038 12.024 7 m/s 4 m/s 6.7 m/s 12.729 3.2731 12.085 7.75 m/s 4 m/s 6.7 m/s 13.469 3.7568 12.596 10 m/s 4 m/s 6.7 m/s 15.031 4.205 14.1 12 m/s 4 m/s 6.7 m/s 16.561 4.655 15.226 Table 6: Measured Data regarding mutations in Shear Rate (1/s) at contrasting outlets corresponding to the alteration in sample fluid velocity flow rate (m/s) with conducting fluid considered was water Sample Inlet flow rate Apex Buffer Inlet flow rate Bottom Buffer Inlet flow rate Apex CTC outlet Shear Rate (1/s) Central WBC outlet Shear Rate (1/s) Bottom Outlet Shear Rate (1/s) 4.75 m/s 4 m/s 6.7 m/s 1.5451 x 10^5 52530 2.1209 x 10^5 5 m/s 4 m/s 6.7 m/s 1.6048 x 10^5 54430 1.9276 x 10^5 5.75 m/s 4 m/s 6.7 m/s 1.4822 x 10^5 55147 2.1039 x 10^5 6 m/s 4 m/s 6.7 m/s 1.6027 x 10^5 55414 2.3421 x 10^5 6.75 m/s 4 m/s 6.7 m/s 1.7987 x 10^5 59386 2.3805 x 10^5 7 m/s 4 m/s 6.7 m/s 2.0410 x 10^5 61118 2.5482 x 10^5 7.75 m/s 4 m/s 6.7 m/s 2.1430 x 10^5 63727 2.5677 x 10^5 10 m/s 4 m/s 6.7 m/s 2.2095 x 10^5 69827 2.8721 x 10^5 12 m/s 4 m/s 6.7 m/s 2.3774 x 10^5 71843 3.018 x 10^5 Table 7: Measured Data regarding mutations in Velocity magnitude (m/s) at contrasting outlets corresponding to the alteration in sample fluid velocity flow rate (m/s) with conducting fluid considered was water Sample Inlet flow rate Apex Buffer Inlet flow rate Bottom Buffer Inlet flow rate Apex CTC outlet velocity magnitude Central WBC outlet velocity magnitude Bottom Outlet velocity magnitude 4.75 m/s 4 m/s 6.7 m/s 9.0016 m/s 2.4161 m/s 8.4055 m/s 5 m/s 4 m/s 6.7 m/s 9.0730 m/s 2.4843 m/s 8.6381 m/s 5.75 m/s 4 m/s 6.7 m/s 9.4907 m/s 2.5244 m/s 8.9178 m/s 6 m/s 4 m/s 6.7 m/s 9.6798 m/s 2.5863 m/s 9.3707 m/s 6.75 m/s 4 m/s 6.7 m/s 10.093 m/s 2.6936 m/s 9.6615 m/s 7 m/s 4 m/s 6.7 m/s 10.357 m/s 2.7360 m/s 9.4361 m/s 7.75 m/s 4 m/s 6.7 m/s 10.596 m/s 2.8310 m/s 9.8434 m/s 10 m/s 4 m/s 6.7 m/s 11.805 m/s 3.1623 m/s 11.178 m/s 12 m/s 4 m/s 6.7 m/s 12.892 m/s 3.5252 m/s 12.096 m/s Figure 14 depicts the mutations in pressure (Pa), Figure 15 depicts the mutations in Reynolds number (Re), Figure 16 depicts the mutations in Shear Rate (1/s) and Figure 17 depicts the mutations in Velocity magnitude (m/s) on the three contrasting outlets corresponding to the alteration in fluid flow rate (m/s) of sample. The graph evidently displays the consistency in pressure (Pa) dissemination, Reynolds number (Re), Shear Rate (1/s) and Velocity magnitude (m/s) at contrasting outlets post the point of separation of Circulating Tumor Cells (CTCs) which was at 12.3 m/s fluid velocity flow rate whenever conducting fluid was used as Blood. Table 8: Measured Data regarding mutations in pressure (Pa) at contrasting outlets corresponding to the alteration in sample fluid velocity flow rate (m/s) with conducting fluid considered was blood Sample Inlet flow rate Apex Buffer Inlet flow rate Bottom Buffer Inlet flow rate Apex CTC outlet Pressure (Pa) Central WBC outlet Pressure (Pa) Bottom Outlet Pressure (Pa) 12.3 m/s 9.5 m/s 15.2 m/s 3475.5 870.54 1392.9 13 m/s 9.5 m/s 15.2 m/s 3578.2 970.71 1594 13.3 m/s 9.5 m/s 15.2 m/s 3827.6 991.95 1602.2 14 m/s 9.5 m/s 15.2 m/s 4158.9 1006.7 1828.8 14.3 m/s 9.5 m/s 15.2 m/s 4427 1028.5 2036.5 15 m/s 9.5 m/s 15.2 m/s 4558.2 1043.8 2697.6 15.3 m/s 9.5 m/s 15.2 m/s 4615 1066.3 3431.5 16 m/s 9.5 m/s 15.2 m/s 4966.2 1082.4 3543.1 16.3 m/s 9.5 m/s 15.2 m/s 5190.1 1105.4 3535.9 Table 9: Measured Data regarding mutations in Reynolds Number (Re) at contrasting outlets corresponding to the alteration in sample fluid velocity flow rate (m/s) with conducting fluid considered was blood Sample Inlet flow rate Apex Buffer Inlet flow rate Bottom Buffer Inlet flow rate Apex CTC outlet Cell Reynolds number Central WBC outlet Cell Reynolds number Bottom Outlet Cell Reynolds number 12.3 m/s 9.5 m/s 15.2 m/s 10.644 2.9015 9.6467 13 m/s 9.5 m/s 15.2 m/s 10.931 2.9908 10.003 13.3 m/s 9.5 m/s 15.2 m/s 10.967 3.0866 10.155 14 m/s 9.5 m/s 15.2 m/s 11.097 3.1324 10.278 14.3 m/s 9.5 m/s 15.2 m/s 11.234 3.319 10.282 15 m/s 9.5 m/s 15.2 m/s 11.866 3.266 10.471 15.3 m/s 9.5 m/s 15.2 m/s 12.097 3.287 10.234 16 m/s 9.5 m/s 15.2 m/s 12.117 3.344 10.645 16.3 m/s 9.5 m/s 15.2 m/s 12.37 3.3677 11.03 Table 10: Measured Data regarding mutations in Shear Rate (1/s) at contrasting outlets corresponding to the alteration in sample fluid velocity flow rate (m/s) with conducting fluid considered was blood Sample Inlet flow rate Apex Buffer Inlet flow rate Bottom Buffer Inlet flow rate Apex CTC outlet Shear Rate (1/s) Central WBC outlet Shear Rate (1/s) Bottom Outlet Shear Rate (1/s) 12.3 m/s 9.5 m/s 15.2 m/s 3.77 x 10^5 1.394 x 10^5 4.98 x 10^5 13 m/s 9.5 m/s 15.2 m/s 4.363 x 10^5 1.4499 x 10^5 5.28 x 10^5 13.3 m/s 9.5 m/s 15.2 m/s 5.664 x 10^5 1.541 x 10^5 5.372 x 10^5 14 m/s 9.5 m/s 15.2 m/s 4.663 x 10^5 1.574 x 10^5 5.3499 x 10^5 14.3 m/s 9.5 m/s 15.2 m/s 5.634 x 10^5 1.667 x 10^5 5.527 x 10^5 15 m/s 9.5 m/s 15.2 m/s 7.067 x 10^5 1.795 x 10^5 5.069 x 10^5 15.3 m/s 9.5 m/s 15.2 m/s 8.087 x 10^5 1.905 x 10^5 5.680 x 10^5 16 m/s 9.5 m/s 15.2 m/s 8.199 x 10^5 1.927 x 10^5 5.443 x 10^5 16.3 m/s 9.5 m/s 15.2 m/s 8.314 x 10^5 2.219 x 10^5 7.724 x 10^5 Table 11: Measured Data regarding mutations in Velocity magnitude (m/s) at contrasting outlets corresponding to the alteration in sample fluid velocity flow rate (m/s) with conducting fluid considered was blood Sample Inlet flow rate Apex Buffer Inlet flow rate Bottom Buffer Inlet flow rate Apex CTC outlet velocity magnitude Central WBC outlet velocity magnitude Bottom Outlet velocity magnitude 12.3 m/s 9.5 m/s 15.2 m/s 21.420 m/s 5.9766 m/s 20.717 m/s 13 m/s 9.5 m/s 15.2 m/s 21.803 m/s 6.0082 m/s 20.636 m/s 13.3 m/s 9.5 m/s 15.2 m/s 21.931 m/s 6.066 m/s 20.819 m/s 14 m/s 9.5 m/s 15.2 m/s 22.468 m/s 6.1925 m/s 20.905 m/s 14.3 m/s 9.5 m/s 15.2 m/s 22.552 m/s 6.294 m/s 21.045 m/s 15 m/s 9.5 m/s 15.2 m/s 22.988 m/s 6.301 m/s 21.591 m/s 15.3 m/s 9.5 m/s 15.2 m/s 23.229 m/s 6.4019 m/s 21.958 m/s 16 m/s 9.5 m/s 15.2 m/s 23.549 m/s 6.4649 m/s 22.117 m/s 16.3 m/s 9.5 m/s 15.2 m/s 23.757 m/s 6.5721 m/s 22.312 m/s Estimation of values for the separation/sorting ratio in percentage was approximated depending on trajectories of bioparticles passing through contrasting outlets and comparable results are shown in Table 12 and Table 13 and the plot shows the diversity of separation/sorting ratio with respect to velocity fluid flow rate and mass fluid flow rate depicted in Figure 18 and Figure 19 respectively. The graph evidently shows that maximal separation/sorting ratio was recognized at a sample velocity fluid flow rate of about 12.3 m/s and sample mass fluid flow rate of 4 x 10-6 kg/s whenever conducting fluid was used as Blood Table 12: Estimated Separation/sorting Ratio (%) of bioparticles/cells viaindividual outlets for distinct fluid velocity flow rates Sample Velocity Fluid Flow rate (m/s) whenever conducting fluid was used as Blood Separation/Sorting ratio (%) of CTCs via apex CTC Outlet (Approx.) Separation/Sorting ratio (%) of WBCs via Central WBC Outlet (Approx.) 11 69 64 11.3 71 62 12 83 79 12.3 87.5 95 13 67.5 70 13.3 65 65 Figure 18apparentlyillustrates the diversity of estimated separation/Sorting ratio of WBCs and CTCs via central outlet for WBC and upper outlet for CTC correspondingly with the adjustment in sample velocity fluid flow rate (m/s). Maximal isolation/sorting ratio for WBCs and CTCs were achieved as 95 % and 87.5 % correspondingly at sample fluid flow rate of 12.3 m/s whenever conducting fluid was used as Blood Table 13: Estimated Separation/sorting Ratio (%) of bioparticles/cells via individual outlets for distinct mass fluid flow rates Sample Mass Fluid Flow rate (kg/s) Separation/Sorting ratio (%) of CTCs via apex CTC Outlet (Approx.) Separation/Sorting ratio (%) of WBCs via Central WBC Outlet (Approx.) 2.5 x 10-6 69 65 3 x 10-6 73 78 3.5 x 10-6 80 92 4 x 10-6 84 96 4.5 x 10-6 82 91 5 x 10-6 72 80 Figure 19 apparently illustrates the diversity of estimated separation/Sorting ratio of WBCs and CTCs via central outlet for WBC and upper outlet for CTC correspondingly with the adjustment in sample mass fluid flow rate (Kg/s). Maximal isolation/sorting ratio for WBCs and CTCs were achieved as 96 % and 84 % correspondingly at sample fluid flow rate of 4 x 10-6 kg/s whenever conducting fluid was used as Blood. CONCLUSION The present invention is focused at designing an effective micro channel device for seclusion of CTCs from WBCs investigates and validates the importance of inclusion of the Deterministic Lateral Displacement principles for accurately planned sorting of blood cells. An efficient micro channel to isolate the CTCs from WBCs from an inflicted blood sample has been modeled using optimized parameters involved in DLD using COMSOL Multiphysics 5.4 software simulation techniques. The proposed geometrical layout of the micro channel device supports both for lower mass flow rates and higher mass flow rates, typically tested at 4 × 10–6 kg/s displaying suitable isolation ratios of 87.5% and 84%, respectively, at 12.5 m/s velocity flow rate and 4 × 10 -6kg/s mass flow rate, making it implementable specially for high-throughput operations. The simulated micro channel device was further tested with velocity flow rate and results indicate that it uniformly isolates CTCs at a high sample velocity flow rate of 12.5 m/s. Moreover, a linear relationship between the variation in critical fluid flow characteristics like pressure, velocity magnitude, shear rate and Reynolds number was established by tapping the measured values from the simulation response as per the alteration in velocity flow rate of the injected sample. In addition to linearity in fluid parameters, the micro channel model exhibits balanced diffusion of fluid pressure inside the network due to symmetricity in alignment of three designated outlets, thereby attenuating risk of clogging. The proposed micro channel device design, therefore offers superior suitability towards seclusion of CTCs from WBCs providing higher isolation efficiency, performing better in high throughput operations, showing linearity of flow characteristics and restricting clogging concerns substantially when compared with presently existing designs in this domain. REFERENCES. Zhou, J., Kulasinghe, A., Bogseth, A. et al., “Isolation of circulating tumor cells in non-small-cell-lung-cancer patients using a multi-flow microfluidic channel”, Microsystems &Nanoengineering, Volume 5, Issue 8, 2019. Jonathan Kottmeier, MaikeWullen weber, Sebastian Blahout, Jeanette Hussong, Ingo Kampen, Arno Kwade and Andreas Dietzel, “Accelerated Particle Separation in a DLD Device at Re > 1 Investigated by Means of µPIV”, Micromachines, Volume 10,pp 768, 2019. 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