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The limitation of fibrous hydrogels to narrow capillaries is of great importance in biological and biomedical systems. Tension and uniaxial compression of fibrous hydrogels have been extensively studied, but their response to biaxial retention in capillaries remains unexplored. Here, we demonstrate experimentally and theoretically that filamentous gels respond qualitatively differently to constraint than flexible chain gels due to the asymmetry in the mechanical properties of the constituent filaments, which are soft in compression and stiff in tension. Under strong retention, the fibrous gel exhibits little elongation and an asymptotic decrease in biaxial Poisson’s ratio to zero, resulting in strong gel compaction and poor liquid permeation through the gel. These results indicate the resistance of stretched occlusive thrombi to lysis by therapeutic agents and stimulate the development of effective endovascular embolization from fibrous gels to stop vascular bleeding or inhibit the blood supply of tumors.
Fibrous networks are the basic structural and functional building blocks of tissues and living cells. Actin is a major component of the cytoskeleton1; fibrin is a key element in wound healing and thrombus formation2, and collagen, elastin and fibronectin are components of the extracellular matrix in the animal kingdom3. Recovered networks of fibrous biopolymers have become materials with wide applications in tissue engineering4.
Filamentous networks represent a separate class of biological soft matter with mechanical properties that are different from flexible molecular networks5. Some of these properties have evolved in the course of evolution to control the response of biological matter to deformation6. For example, fibrous networks show linear elasticity at small strains7,8 while at large strains they exhibit increased stiffness9,10, thereby maintaining tissue integrity. The implications for other mechanical properties of fibrous gels, such as negative normal stress in response to shear strain11,12, have yet to be discovered.
The mechanical properties of semi-flexible fibrous hydrogels have been studied under uniaxial tension13,14 and compression8,15, but their freedom-induced biaxial compression in narrow capillaries or tubes has not been studied. Here we report experimental results and theoretically propose a mechanism for the behavior of fibrous hydrogels under biaxial retention in microfluidic channels.
Fibrin microgels with various ratios of fibrinogen and thrombin concentrations and a D0 diameter ranging from 150 to 220 µm were generated using a microfluidic approach (Supplementary Figure 1). On fig. 1a shows images of fluorochrome labeled microgels obtained using confocal fluorescence microscopy (CFM). The microgels are spherical, have a polydispersity of less than 5%, and are uniform in structure across the scales examined by CFM (Supplementary Information and Movies S1 and S2). The average pore size of the microgels (determined by measuring the Darcy permeability16) decreased from 2280 to 60 nm, the fibrin content increased from 5.25 to 37.9 mg/mL, and the thrombin concentration decreased from 2.56 to 0.27 units/mL, respectively. (Additional Information). Rice. 2), 3 and supplementary table 1). The corresponding stiffness of the microgel increases from 0.85 to 3.6 kPa (Supplementary Fig. 4). As examples of gels formed from flexible chains, agarose microgels of various stiffnesses are used.
Fluorescence microscopy image of fluorescein isothiocyanate (FITC) labeled PM suspended in TBS. The bar scale is 500 µm. b SEM images of SM (top) and RM (bottom). Scale bar 500 nm. c Schematic diagram of a microfluidic channel consisting of a large channel (diameter dl) and a narrowed cone-shaped region with an entry angle α of 15° and a diameter of dc = 65 µm. d Left to right: Optical microscope images of RM (diameter D0) in large channels, conical zone and constriction (limiting gel length Dz). The bar scale is 100 µm. e, f TEM images of an undeformed RM (e) and an occluded RM (f), fixed for one hour with constriction 1/λr = 2.7, followed by release and fixation of 5% of the mass. glutaraldehyde in TBS. The diameter of the undeformed CO is 176 μm. The scale bar is 100 nm.
We focused on fibrin microgels with a hardness of 0.85, 1.87 and 3.6 kPa (hereinafter referred to as soft microgels (SM), medium hard microgels (MM) and hard microgels (RM), respectively). This range of fibrin gel stiffness is of the same order of magnitude as for blood clots18,19 and hence the fibrin gels studied in our work are directly related to real biological systems. On fig. 1b shows the top and bottom images of the SM and RM structures obtained using a scanning electron microscope (SEM), respectively. Compared to RM structures, SM networks are formed by thicker fibers and fewer branch points, consistent with earlier reports 20, 21 (Supplementary Fig. 5). The difference in the structure of the hydrogel correlates with the trend of its properties: the permeability of the gel decreases with decreasing pore size from SM to MM and RM (Supplementary Table 1), and the stiffness of the gel reverses. No changes in microgel structure were noted after storage at 4 °C for 30 days (Supplementary Fig. 6).
On fig. 1c shows a diagram of a microfluidic channel with a circular cross section containing (from left to right): a large channel with a diameter dl in which the microgel remains undeformed, a cone-shaped section with a narrowing in diameter dc < D0, cone-shaped sections and large channels with a diameter dl (Supplementary Fig. 7). In a typical experiment, microgels were injected into microfluidic channels at a positive pressure drop ΔP of 0.2–16 kPa (Supplementary Fig. 8). This pressure range corresponds to biologically significant blood pressure (120 mm Hg = 16 kPa)22. On fig. 1d (from left to right) shows representative images of RM in large channels, conical areas and constrictions. The movement and shape of the microgel were recorded and analyzed using the MATLAB program. It is important to note that in the tapering regions and constrictions, the microgels are in conformal contact with the walls of the microchannels (Supplementary Fig. 8). The degree of radial retention of the microgel at narrowing D0/dc = 1/λr is in the range 2.4 ≤ 1/λr ≤ 4.2, where 1/λr is the compression ratio. The microgel goes through shrinkage when ΔP > ΔPtr, where ΔPtr is the translocation pressure difference. The length and size of the pores of biaxially constrained microgels are determined by their equilibrium state, since it is very important to take into account the viscoelasticity of gels in biological systems. The equilibration time for agarose and fibrin microgels was 10 min and 30 min, respectively. After these time intervals, the limited microgels reached their stable position and shape, which was captured using a high-speed camera and analyzed using MATLAB.
On fig. 1e, 1f show transmission electron microscopy (TEM) images of undeformed and biaxially limited RM structures. After RM compression, the microgel pore size significantly decreased and their shape became anisotropic with smaller sizes in the direction of compression, which is consistent with an earlier report 23 .
Biaxial compression during contraction causes the microgel to elongate in an unlimited direction with a coefficient λz = \({D}_{{{{{{{\rm{z}}}}}}}/\({D }_ { 0}\) , where \({D}_{{{{({\rm{z}}}}}}}}\) is the length of the closed microgel Figure 2a shows the change in λzvs .1/ λr for fibrin and agarose microgels. Surprisingly, under strong compression of 2.4 ≤ 1/λr ≤ 4.2, fibrin microgels show a negligible elongation of 1.12 +/- 0.03 λz, which is only slightly affected by the value of 1/λr. the behavior of limited agarose microgels, which are observed even at weaker compression 1/λr = 2.6 to a larger elongation λz = 1.3.
a Agarose microgel experiments with different elastic moduli (2.6 kPa, green open diamond; 8.3 kPa, brown open circle; 12.5 kPa, orange open square; 20.2 kPa, magenta open inverted triangle) and SM ( solid red) Change in measured elongation λz (circles), MM (solid black squares) and RM (solid blue triangles). Solid lines show the theoretically predicted λz for agarose (green line) and fibrin microgels (lines and symbols of the same color). b, c Top panel: schematic diagram of network chains of agarose (b) and fibrin (c) before (left) and after (right) biaxial compression. Bottom: Shape of the corresponding network before and after deformation. The x and y compression directions are indicated by magenta and brown arrows, respectively. In the figure above, chains of networks oriented in these x and y directions are shown with the corresponding magenta and brown lines, and chains oriented in an arbitrary z direction are represented by green lines. In the fibrin gel (c), the purple and brown lines in the x and y directions bend more than in the undeformed state, and the green lines in the z direction bend and stretch. The tension between the directions of compression and tension is transmitted through threads with intermediate directions. In agarose gels, the chains in all directions determine the osmotic pressure, which makes a significant contribution to the deformation of the gel. d Predicted change in biaxial Poisson’s ratio, } }^{{{{{\rm{eff}}}}}}} =-{{{{{\rm{ln}}}}}}{\lambda }_{z}/{{{{{ {{ \rm{ln}}}}}}{\lambda }_{r}\ ), for equibiaxial compression of agarose (green line) and fibrin (red line) gels. The inset shows the biaxial deformation of the gel. e Translocation pressure change ΔPtr, normalized to gel stiffness S, is plotted as a function of compression ratio for agarose and fibrin microgels. The symbol colors correspond to the colors in (a). The green and red lines depict the theoretical relationship between ΔPtr/S and 1/λr for agarose and fibrin gels, respectively. The dashed part of the red line shows the increase in ΔPtr under strong compression due to interfiber interactions.
This difference is associated with different mechanisms of deformation of fibrin and agarose microgel networks, which consist of flexible24 and rigid25 threads, respectively. Biaxial compression of flexible gels leads to a decrease in their volume and an associated increase in concentration and osmotic pressure, which leads to an elongation of the gel in an unlimited direction. The final elongation of the gel depends on the balance of an increase in the entropic free energy of the stretched chains and a decrease in the free energy of osmosis due to the lower polymer concentration in the stretched gel. Under strong biaxial compression, the elongation of the gel increases with λz ≈ 0.6 \({{\lambda}_{{{\rm{r}}}}^{-2/3}}\) (see Fig. 2a in discussion section 5.3.3). The conformational changes in flexible chains and the shape of the corresponding networks before and after biaxial retention are shown in Figs. 2b.
In contrast, fibrous gels such as fibrin inherently respond differently to biaxial retention. The filaments oriented predominantly parallel to the direction of compression flex (thereby reducing the distance between the cross-links), while the filaments predominantly perpendicular to the direction of compression straighten and stretch under the action of the elastic force, causing the gel to elongate (Fig. 1). 2c) The structures of the undeformed SM, MM and RM were characterized by analyzing their SEM and CFM images (Supplementary Discussion Section IV and Supplementary Figure 9). By determining the elastic modulus (E), diameter (d), profile length (R0), distance between ends (L0 ≈ R0) and central angle (ψ0) of the strands in undeformed fibrin microgels (Supplementary Table 2) – 4), we find that thread bending modulus \({k}_{{{{{{\rm{b)))))))))}=\frac{9\pi E{d}^{4} } {4{\psi } _{0}^{2}{L}_{0}}\) is significantly less than its tensile modulus\({k}_{{{{{{{\rm{s}}} } }} }}=E\frac{\pi {d}^{2}{R}_{0}}{4}\), so kb/ks ≈ 0.1 (Supplementary Table 4). Thus, under conditions of biaxial gel retention, fibrin strands are easily bent, but resist stretching. The elongation of a filamentous network subjected to biaxial compression is shown in Supplementary Fig. 17.
We develop a theoretical affine model (Supplementary Discussion Section V and Supplementary Figures 10–16) in which the elongation of a fibrous gel is determined from the local equilibrium of the elastic forces acting in the gel and predicts that in a strong biaxial strain λz -1 under the constraint
Equation (1) shows that even under strong compression (\({\lambda }_{{{\mbox{r))))\,\to \,0\)) there is a slight gel expansion and subsequent elongation deformation upon saturation λz–1 = 0.15 ± 0.05. This behavior is related to (i) \({\left({k}_{{{{({\rm{b}}}}}}}}}/{k}_{{{{{{\rm{ s }}}}}}}\right)}^{1/2}\) ≈ 0.15−0.4 and (ii) the term in square brackets asymptotically approximates \(1{{\mbox{/}}} \sqrt{ 3 }\) for strong biaxial bonds. It is important to note that the prefactor \({\left({k}_{({\mbox{b))))/{k}_{({\mbox{s))))\right)}^{1/ 2 }\) has nothing to do with the stiffness of the thread E, but is determined only by the aspect ratio of the thread d/L0 and the central angle of the arc ψ0, which is similar to SM, MM and RM (Supplementary Table 4).
To further highlight the difference in freedom-induced strain between flexible and filamentous gels, we introduce the biaxial Poisson’s ratio \({\nu }_{{{({\rm{b)))))) }{{\ mbox { =}}}\,\mathop{{\lim}}\limits_{{\lambda}_{{{{({\rm{r}}}}}}}\to 1}\ frac{{\ lambda }_{ {{{{\rm{z}}}}}}-1}{1-{\lambda }_{{({\rm{r}}}}}}}}}, \) describes an unbounded orientation of gel strain in response to equal strain in two radial directions, and extends this to large uniform strains \ rm{b }}}}}}}}^{{{{{\rm{eff}}}}}}}}}=-{{{{{\rm{ln}}}}}}} }{ \lambda } _{z} /{{{({\rm{ln)))))))}{\lambda }_{{{({\rm{r)))))))))}\) . On fig. 2d shows \({{{{{{\rm{\nu }}}}}}}_{{{({\rm{b}}}}}}}}^{{{ {{\rm { eff}}}}}}}\) for uniform biaxial compression of flexible (such as agarose) and rigid (such as fibrin) gels (Supplementary discussion, Section 5.3.4), and highlights the relationship between strong differences in responses to confinement. For agarose gels under strong restrictions {\rm{eff}}}}}}}}\) increases to the asymptotic value 2/3, and for fibrin gels it decreases to zero, since lnλz/lnλr → 0, since λz increases with saturation as λr increases. Note that in experiments, closed spherical microgels deform inhomogeneously, and their central part experiences stronger compression; however, extrapolation to a large value of 1/λr makes it possible to compare the experiment with the theory for uniformly deformed gels.
Another difference in the behavior of flexible chain gels and filamentous gels was found due to their movement upon contraction. The translocation pressure ΔPtr, normalized to gel stiffness S, increased with increasing compression (Fig. 2e), but at 2.0 ≤ 1/λr ≤ 3.5, fibrin microgels showed significantly lower values of ΔPtr/S down during shrinkage. Retention of the agarose microgel leads to an increase in osmotic pressure, which leads to the stretching of the gel in the longitudinal direction as the polymer molecules are stretched (Fig. 2b, left) and an increase in translocation pressure by ΔPtr/S ~(1/λr)14/317. On the contrary, the shape of closed fibrin microgels is determined by the energy balance of the threads of radial compression and longitudinal tension, which leads to the maximum longitudinal deformation λz ~\(\sqrt{{k}_{{{ {{ {\rm{ b)))))))} /{k}_{{{{{{{\rm{s}}}}}}}}}\). For 1/λr ≫ 1, the change in translocation pressure is scaled as 1 }{{{({\rm{ln))))))\left({{\lambda }}_{{{{{{\rm{r} }}}}}}}^{{-} 1} \right)\) (Supplementary Discussion, Section 5.4), as shown by the solid red line in Fig. 2e. Thus, ΔPtr is less constrained than in agarose gels. For compressions with 1/λr > 3.5, a significant increase in the volume fraction of filaments and the interaction of neighboring filaments limits further deformation of the gel and leads to deviations of experimental results from predictions (red dotted line in Fig. 2e). We conclude that for the same 1/λr and Δ\({P}_{{{{{{{\rm{tr}}}}}}}}_{{{{\rm{fibrin}}} ))} }}}\) < ΔP < Δ\({P}_{{{{{{{\rm{tr))))))}}}_{{{{\rm{agarose}} }} }} } }}\) the agarose gel will be captured by the microchannel, and the fibrin gel with the same stiffness will pass through it. For ΔP < Δ\({P}_{{{{{{\rm{tr)))))))))_{{{{{\rm{fibrin)))))))))}\), Two Both gels will block the channel, but the fibrin gel will push deeper and compress more effectively, blocking fluid flow more effectively. The results shown in Figure 2 demonstrate that the fibrous gel can serve as an effective plug to reduce bleeding or inhibit the blood supply to tumors.
On the other hand, fibrin forms a clot scaffold that leads to thromboembolism, a pathological condition in which a thrombus occludes a vessel at ΔP < ΔPtr, such as in some types of ischemic stroke (Fig. 3a). The weaker restriction-induced elongation of fibrin microgels resulted in a stronger increase in fibrin concentration of C/C fibrinogen compared to flexible chain gels, where C and C fibrinogen are restricted and undeformed microgels, respectively. Polymer concentration in the gel. Figure 3b shows that fibrinogen C/C in SM, MM, and RM increased more than seven-fold at 1/λr ≈ 4.0, driven by restriction and dehydration (Supplementary Fig. 16).
Schematic illustration of occlusion of the middle cerebral artery in the brain. b Restriction-mediated relative increase in fibrin concentration in obstructive SM (solid red circles), MM (solid black squares), and RM (solid blue triangles). c Experimental design used to study the cleavage of restricted fibrin gels. A solution of fluorescently labeled tPA in TBS was injected at a flow rate of 5.6 × 107 µm3/s and an additional pressure drop of 0.7 Pa for channels located perpendicular to the long axis of the main microchannel. d Pooled multichannel microscopic image of obstructive MM (D0 = 200 µm) at Xf = 28 µm, ΔP = 700 Pa and during splitting. Vertical dotted lines show the initial positions of the posterior and anterior edges of the MM at tlys = 0. Green and pink colors correspond to FITC-dextran (70 kDa) and tPA labeled with AlexaFluor633, respectively. e Time-varying relative volume of occluded RMs with D0 of 174 µm (blue open inverted triangle), 199 µm (blue open triangle), and 218 µm (blue open triangle), respectively, in a conical microchannel with Xf = 28 ± 1 µm. the sections have ΔP 1200, 1800, and 3000 Pa, respectively, and Q = 1860 ± 70 µm3/s. The inset shows RM (D0 = 218 µm) plugging the microchannel. f Time variation of the relative volume of SM, MM or RM placed at Xf = 32 ± 12 µm, at ΔP 400, 750 and 1800 Pa and ΔP 12300 Pa and Q 12300 in the conical region of the microchannel, respectively 2400 and 1860 µm3/s. Xf represents the front position of the microgel and determines its distance from the start of shrinkage. V(tlys) and V0 are the temporary volume of the lysed microgel and the volume of the undisturbed microgel, respectively. The character colors correspond to the colors in b. Black arrows on e, f correspond to the last moment of time before the passage of microgels through the microchannel. The scale bar in d, e is 100 µm.
To investigate the effect of restriction on fluid flow reduction across obstructive fibrin gels, we studied the lysis of SM, MM, and RM infiltrated with the thrombolytic agent tissue plasminogen activator (tPA). Figure 3c shows the experimental design used for the lysis experiments. At ΔP = 700 Pa (<ΔPtr) and a flow rate, Q = 2400 μm3/s, of Tris-buffered saline (TBS) mixed with 0.1 mg/mL of (fluorescein isothiocyanate) FITC-Dextran, the microgel occluded the tapered microchannel region. At ΔP = 700 Pa (<ΔPtr) and a flow rate, Q = 2400 μm3/s, of Tris-buffered saline (TBS) mixed with 0.1 mg/mL of (fluorescein isothiocyanate) FITC-Dextran, the microgel occluded the tapered microchannel region. При ΔP = 700 Па (<ΔPtr) и скорости потока, Q = 2400 мкм3/с, трис-буферного солевого раствора (TBS), смешанного с 0,1 мг/мл (флуоресцеинизотиоцианата) FITC-декстрана, микрогель перекрывал сужающийся микроканал. At ΔP = 700 Pa (<ΔPtr) and a flow rate, Q = 2400 µm3/s, of Tris buffered saline (TBS) mixed with 0.1 mg/mL (fluorescein isothiocyanate) FITC-dextran, the microgel occluded the converging microchannel. region.在ΔP = 700 Pa (<ΔPtr) 和流速Q = 2400 μm3/s 的Tris 缓冲盐水(TBS) 与0.1 mg/mL 的（异硫氰酸荧光素）FITC-葡聚糖混合时，微凝胶堵塞了锥形微通道地区。在ΔP = 700 Pa (<ΔPtr) 和流速Q = 2400 μm3/s了锥形微通道地区。 Микрогели закупориваются при смешивании трис-буферного солевого раствора (TBS) с 0,1 мг/мл (флуоресцеинизотиоцианат) FITC-декстрана при ΔP = 700 Па (<ΔPtr) и скорости потока Q = 2400 мкм3/с Конические области микроканалов. Microgels plugged when Tris buffered saline (TBS) was mixed with 0.1mg/mL (fluorescein isothiocyanate) FITC-dextran at ΔP = 700 Pa (<ΔPtr) and flow rate Q = 2400 µm3/s Conical regions of microchannels. The forward position Xf of the microgel determines its distance from the initial shrinkage point X0. To induce lysis, a solution of fluorescently labeled tPA in TBS was injected from a channel located orthogonally to the long axis of the main microchannel.
When the tPA solution reached the occlusal MM, the posterior edge of the microgel became blurred, indicating that fibrin cleavage had begun at time tlys = 0 (Fig. 3d and Supplementary Fig. 18). During fibrinolysis, dye-labeled tPA accumulates inside the MM and binds to fibrin strands, which leads to a gradual increase in the intensity of the pink color of the microgels. At tlys = 60 min, the MM contracts due to the dissolution of its rear part, and the position of its leading edge Xf changes little. After 160 min, the strongly contracted MM continued to contract, and at tlys = 161 min, it underwent contraction, thereby restoring fluid flow through the microchannel (Fig. 3d and Supplementary Fig. 18, right column).
On fig. 3e shows the lysis-mediated time-dependent decrease in volume V(tlys) normalized to the initial volume V0 of different sized fibrin microgels. CO with D0 174, 199, or 218 µm was placed into a microchannel with ΔP 1200, 1800, or 3000 Pa, respectively, and Q = 1860 ± 70 µm3/s to block the microchannel (Fig. 3e, inset). nutrition. The microgels gradually shrink until they are small enough to pass through the channels. A decrease in the critical volume of CO with a larger initial diameter requires a longer lysis time. Due to the similar flow through different sized RMs, cleavage occurs at the same rate, resulting in digestion of smaller fractions of larger RMs and their delayed translocation. On fig. 3f shows the relative reduction in V(tlys)/V0 due to splitting for SM, MM, and RM at D0 = 197 ± 3 µm plotted as a function of tlys. For SM, MM and RM, place each microgel in a microchannel with ΔP 400, 750 or 1800 Pa and Q 12300, 2400 or 1860 µm3/s, respectively. Although the pressure applied to the SM was 4.5 times lower than that of the RM, the flow through the SM was more than six times stronger due to the higher permeability of the SM, and the shrinkage of the microgel decreased from SM to MM and RM. For example, at tlys = 78 min, SM mostly dissolved and displaced, while MM and PM continued to clog the microchannels, despite retaining only 16% and 20% of their original volume, respectively. These results suggest the importance of convection-mediated lysis of constricted fibrous gels and correlate with reports of faster digestion of clots with lower fibrin content.
Thus, our work demonstrates experimentally and theoretically the mechanism by which filamentous gels respond to biaxial confinement. The behavior of fibrous gels in a limited space is determined by the strong asymmetry of the strain energy of the filaments (soft in compression and hard in tension) and only by the aspect ratio and curvature of the filaments. This reaction results in minimal elongation of fibrous gels contained in narrow capillaries, their biaxial Poisson’s ratio decreasing with increasing compression and less light bit pressure.
Since biaxial containment of soft deformable particles is used in a wide range of technologies, our results stimulate the development of new fibrous materials. In particular, the biaxial retention of filamentous gels in narrow capillaries or tubes leads to their strong compaction and a sharp decrease in permeability. The strong inhibition of fluid flow through occlusive fibrous gels has advantages when used as plugs to prevent bleeding or reduce the blood supply to malignancies33,34,35. On the other hand, a decrease in fluid flow through the occlusal fibrin gel, thereby inhibiting convective-mediated thrombus lysis, gives an indication of the slow lysis of occlusal clots [27, 36, 37]. Our modeling system is the first step towards understanding the implications of the mechanical response of fibrous biopolymer hydrogels to biaxial retention. Incorporating blood cells or platelets into obstructive fibrin gels will affect their restriction behavior 38 and will be the next step in uncovering the behavior of more complex biologically significant systems.
Reagents used to prepare fibrin microgels and fabricate MF devices are described in Supplementary Information (Supplementary Methods Sections 2 and 4). Fibrin microgels were prepared by emulsifying a mixed solution of fibrinogen, Tris buffer and thrombin in a flow focusing MF device, followed by droplet gelation. Bovine fibrinogen solution (60 mg/ml in TBS), Tris buffer and bovine thrombin solution (5 U/ml in 10 mM CaCl2 solution) were administered using two independently controlled syringe pumps (PhD 200 Harvard Apparatus PHD 2000 Syring Pump). to block MF, USA). F-oil continuous phase containing 1 wt.% block copolymer PFPE-P(EO-PO)-PFPE, was introduced into the MF unit using a third syringe pump. Droplets formed in the MF device are collected in a 15 ml centrifuge tube containing F-oil. Place the tubes in a water bath at 37 °C for 1 h to complete fibrin gelation. FITC labeled fibrin microgels were prepared by mixing bovine fibrinogen and FITC labeled human fibrinogen in a 33:1 weight ratio, respectively. The procedure is the same as for the preparation of fibrin microgels.
Transfer the microgels from oil F to TBS by centrifuging the dispersion at 185 g for 2 min. The precipitated microgels were dispersed in oil F mixed with 20 wt.% perfluorooctyl alcohol, then dispersed in hexane containing 0.5 wt.% Span 80, hexane, 0.1 wt.% Triton X in water and TBS. Finally, the microgels were dispersed in TBS containing 0.01 wt% Tween 20 and stored at 4°C for approximately 1–2 weeks prior to experiments.
The fabrication of the MF device is described in the Supplementary Information (Supplementary Methods Section 5). In a typical experiment, the positive value of ΔP is determined by the relative height of the reservoirs connected before and after the MF device for introducing microgels with a diameter of 150 < D0 < 270 µm into the microchannels. The undisturbed size of the microgels was determined by visualizing them in the macrochannel. The microgel stops in a conical area at the entrance to the constriction. When the tip of the anterior microgel remains unchanged for 2 min, use the MATLAB program to determine the position of the microgel along the x-axis. With a stepwise increase in ΔP, the microgel moves along the wedge-shaped region until it enters the constriction. Once the microgel is fully inserted and compressed, ΔP rapidly drops to zero, balancing the water level between the reservoirs, and the closed microgel remains stationary under compression. The length of the obstructive microgel was measured 30 min after the constriction ceased.
During fibrinolysis experiments, solutions of t-PA and FITC-labeled dextran penetrate blocked microgels. The flow of each liquid was monitored using single channel fluorescence imaging. TAP labeled with AlexaFluor 633 attached to fibrin fibers and accumulated inside compressed fibrin microgels (TRITC channel in Supplementary Fig. 18). The dextran solution labeled with FITC moves without accumulation in the microgel.
Data supporting the results of this study are available from the respective authors upon request. Raw SEM images of fibrin gels, raw TEM images of fibrin gels before and after inoculation, and the main input data for Figures 1 and 2. 2 and 3 are provided in the raw data file. This article provides the original data.
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