WO2022035335A1 - Fabrication of microactuator and in-line degasser in organ-ona_chip devices and methods thereof - Google Patents

Abstract

The present invention refers to a method for the fabrication of complex organ-on-a-chip devices with incorporated mechanical microactuator and bubble degasser. The method herein described is based on the micro-structuring of polydimethylsiloxane (PDMS) laminates, using a cutting plotter to generate multi-layered, membrane integrated biochips in a matter of hours, using low-cost benchtop equipment. Furthermore, a material permeable to gas exchange placed directly underneath the flow chamber and a degassing step allows to operate an integrated bubble degasser. The present invention also describes examples of a device for cell culture of gastric fibroblasts and an epithelial gastric cell line with no cytotoxic effects or impact on cellular homeostasis, stepping towards an advanced model of the stomach mucosa. The devices and methods herein described can be advantageously used in the early stages of organ-on-a-chip development by reducing three of the main burdens facing a microfabrication laboratory, namely production time, cost and space requirement. Fabricated devices retain several desirable qualities such as transparency, flexibility for mechanical actuation, biocompatibility and capability of gaseous exchanges with the exterior environment, allowing elimination of air bubbles formed during operation, which is one of the major challenges during biochip handling.

Description

DESCRIPTION

FABRICATION OF MICROACTUATOR AND IN-LINE DEGASSER IN ORGAN-ON -A-CHIPS DEVICES AND METHODS THEREOF

Technical field of the invention

The present invention relates to the technical field of supports for cultivation of cell lines and tissues characterized by topography and properties .

State of the art

Microfluidic devices are well established as experimental platforms in life sciences (1, 2) . These systems provide the ability to control aspects of the cell microenvironment at more relevant kinetic and spatial scales (3) . The design versatility and scalability of microfluidic platforms, through parallelization of devices in array-like designs (4- 6) , make them a powerful tool. Development of this technology was aided by the advent of soft lithography (7) , a term that refers to a subset of fabrication techniques using the flexible silicon elastomer, polydimethylsiloxane (PDMS) , as the primary building material. The combination of photolithography (8) with the ability of PDMS to faithfully replicate the molded structures at nano level resolution, means that high fidelity, fully customizable devices can be produced in a matter of days, at a relatively low cost (7) . Despite the obvious advantages, soft lithography also has its drawbacks, most notably the techniques employed require access to a clean room facility. Not surprisingly, efforts have been made towards the development of fabrication techniques relying in simpler and space-saving techniques (9) . Alternative methods include milling, laser engraving and paper- based microfluidics (10-12) . This is particularly important in an academic or pre-industrial context, where prototype design must go through several iterations and plastic injection molding is not a viable financial option. 3D printing of thermoplastics has also been explored as it can generate one-piece, enclosed devices, which are of particular interest for cell culture applications (13, 14) .

Culturing cells within these devices requires extensive optimization, despite their many advantages over conventional cell culture (15) . The substantially lower culture chamber volume, the shear stress imparted by the flow of culture media over the cell culture surface, and their relatively low permeability to atmospheric gases , require a careful optimization to ensure an efficient oxygen and nutrient supply and a stable pH throughout the experimental run . This is particularly important when developing organ-on-chip devices . Organ-on-a-chip devices are more complex than cell-based lab-on-a-chip devices , as they aim to recreate the complex micro-phys iological architecture and function of the organ they intend to emulate ( 16-18 ) . Harness ing the cellular microenvironment at the tissue level also requires tackling the impact of mechanical forces at the cellular level and understanding how cells transduce these mechanical forces into biochemical s ignals , in a physiologically relevant context ( 19-24 ) . The success ful integration of microactuators within organ-on-a-chip devices allowed the application of well-defined and cyclical strain on the cell culture substrate . The ability to control the intensity, duration and pattern of the mechanical forces within the system, make organ-on-a-chip platforms a powerful tool to understand how mechanical transduction affects cellular response at the tis sue level , thus modulating to a greater extent a key aspect of the in-vivo native microenvironment . Although ideally suited for the aforementioned tas ks , design of organ-on-chip microdevices is a complex procedure . Their three- dimensional ( 3D) layout and incorporation of several tissue-like features , such as a simplified stromal component and epithelial barrier architecture , is further complicated by the incorporation of embedded mechanical microactuators ( 19-22 ) . Creating complex structures that correctly emulate the biological counterparts , usually implies fabricating multi-layered devices , with two or more chambers separated by porous membranes , as well as complex flexible mechanisms that serve as mechanical actuators . Also , replicating some organ functions requires the des ign of intricate microchannel geometries to house , in a specific layout , the individual components of the organotypic unit being emulated ( 25 , 26 ) or to manipulate diffusion distances ( 5 , 27 ) . While geometry design, by itself , is not constrained by fabrication limitations , it is es sential that they are amenable to correct cell culture maintenance and homeostasis ( 28 ) . Methods with a fast turnaround time , from des ign to device , are key to reduce experimental costs at the early stages of organ-on-a-chip des ign . Additionally, a common complication observed during microfluidic operation is the formation of air bubbles ( 29 ) . Biochips that need to be operated at 37 ° C , are particularly susceptible to this problem as the higher temperature leads to decreased gas solubility ( 2 9 ) . Under these circumstances , the volume of the air bubble will gradually expand over time and air bubbles can disrupt flow and the creation of an air-liquid interface may compromise cellular homeostasis and promote cellular death.

Detailed description of the Invention

The invention discloses a facile fabrication method that is time-saving and permits the fabrication of the described devices at a fraction of the time and cost, when compared to current state of the art technology. Said devices allow application of differential patterns of mechanical actuation on cell culture substrates and can be irreversibly or reversibly bonded to the cell culture fluidic chamber for added versatility. The same cell culture substrate can be subjected to different strain patterns by application of distinct actuator chambers beneath it. Furthermore, a material permeable to gas exchange placed directly underneath the flow chamber and a degassing step allows to operate an integrated bubble degasser.

For microfluidic device layout, the device is designed using a computer aided design (CAD) software and made to fit a 26 x 26 mm plate, suitable for bonding to a standard microscope glass slide (Figure 1) . Fluidic features are rectangular shaped, 2 mm (width) x 10 mm (length) channels, for a total cell culture area of 0.2 cm². Access to the cell culture area is made through a linear feature connecting to an access porthole (Figure 1a-c) . The actuation and degassing chamber (Figure le) is 3.5 mm (width) x 11 mm (length) , topped by a flexible PDMS membrane (Figure Id) . In one embodiment, a full device consists of 9 superimposed layers (Figure 2) . A fluidic upper portion, where liquid and cell handling is performed, contains a top layer of cast-on PDMS, (Sylgard 184, Dow Corning; mixed on a weight ratio of 10:1 PDMS base to curing agent respectively) , where access portholes are punched with a 20 Ga puncher. This is followed by 5 alternating layers of 500 μm PDMS laminate (MVQ silicones) and a cell culture treated polyethylene terephtalate (PET) perforated membrane (thickness of 16 μm, pore size of 8μm, pore density of 6e⁴ cm²) . The bottom part of the device contains the pneumatic actuation and/or the degassing system, consisting of 2, 250 μm PDMS laminate layers (MVQ silicones) placed on top of a glass slide.

For preparation of chip-related materials, prior to assembly, each plate design, is cut from PDMS foil (915 mm wide; MVQ silicones) , using a desktop cutter (model CAMM-1 GS-24; Roland DG) . The fluidic plates are cut from 500 μm thick foil. The plates forming the pneumatic portion of the device are cut from 250 μm thick foil. In one embodiment, a section of PDMS foil of about 3 x 20 cm, is manually cut with scissors and one side of the protective backing plastic is removed. Finally, the laminate sheet is fed to the cutting plotter and the design transferred from theCAD software to a dedicated software, Roland Cut Studio (Roland DG) . After the machining process, the off cut material is removed with tweezers, to reveal the hollow fluidic and pneumatic features. In one embodiment, a PET membrane (thickness of 16 μm, pore size of 8 μm, pore density of 6e4 cm²; it4ip) , is cut manually from 25 mm diameter discs immersed in isopropanol (IPA) , cleaned by ultrasound (5 min) , dried with compressed air and stored in a dust-free container. In one embodiment, the top layer of each chip is fabricated by gravity casting PDMS on a circular petri dish. The PDMS base is mixed thoroughly with curing agent on a weight ratio of PDMS base to curing agent 10:1 (Sylgard 184, Dow Corning) . The heavily aerated pre-polymer mix is degassed by centrifugation (5 min at 4000 rpm) , poured carefully onto 90 mm petri dish and cured at 70 °C for 1 h. The resulting PDMS plate is then cut in 26 mm square sections to match the machined PDMS layers. Glass microscope slides are thoroughly cleaned by sequential incubation (5 min by ultrasound) in 2% (v/v) Helmanex III solution in ddfW (Helma Analytics) , followed by acetone and a final rinse in double distilled H2O . Clean slides are air dried with compressed air and stored in a dust- free container.

For assembly (Figure 2) , each laminate plate is cut individually, with a desktop cutter (CAMM-1 GS-24; Roland DG) and thoroughly cleaned with IPA followed by a drying step with compressed air at 1.5 bar in a dust free environment. PET membranes are cut manually and cleaned in the same manner. PET membranes must first be silanized prior to bonding to ensure long term adhesion during incubation at 37 °C. Briefly, membranes are exposed to O2 plasma (Zepto Plasma Laboratory Unit, Diener) , for 1 min. These are then heated for 20 min in 2% bis [3- ( trimethoxysilyl ) propyl ] amine 1% double distilled H2O in IPA at 80 °C, followed by a rinse with IPA and cured for 30 min at 70 °C. Membranes are then wetted for 30 min in 70% (v/v) ethanol and brought immediately into conformal contact with the PDMS plates previously exposed to 02 plasma. The final structure is obtained by repeating the previous steps, bringing into contact each PDMS plate and PET membrane as needed for each microdevice design (2, 3, 4, etc, layered devices) . The bottom structure of the device is a glass slide, bonded permanently to the PDMS assembly, by exposure to O2 plasma. The whole structure is further baked at 70°C for 1 h.

Microdevice is then primed before cell seeding. Device is initially sterilized by exposure to UV light for 20 min, followed by perfusion of a 70% (v/v) ethanol solution for another twenty minutes. The device is then rinsed 3 times with phosphate buffered saline (PBS) and coated with a fibronectin solution (50μg/mL; Sigma) for 2 h.

The priming step further includes an in-line air bubble degassing (Figure 3) . For this purpose, the chip is filled with PBS solution and allowed to equilibrate at 37°C for at least 2 h. To operate as bubble degasser, a PDMS membrane, which is permeable to gas exchange, is placed directly underneath the flow chamber, spanning its surface, and a constant vacuum pressure of -50 mbar is applied over a period of 4 to 16 h, with the system pressurized and fully contained throughout operation. This is able to effectively eliminate air bubbles over-time from the main culture chamber (Figure 3a) . A -50 mbar constant vacuum pressure can effectively eliminate a bubble with an average area of 250.5 mm² over a period of 4 h (Figure 3b) , which solves a common problem with standard microfluidic systems .

For device seeding, cells maintained in CO2 independent medium (Gibco) , supplemented with 10% (v/v) fetal bovine serum (FBS; Lonza) , 1% (v/v) penicilin/streptomycin (gibco) and 4 mM L-glutamine (Gibco)are employed. In one embodiment, gastric fibroblasts are seeded at a density of 200,000 cells/mL, on the middle perfusion channel, embedded in a collagen type I matrix (3.0 mg/mL; IBIDI) . Collagen type I is allowed to gel for 30 minutes at 37 °C after which the device is coupled to a piezoelectric perfusion control system, OBI (Elveflow) . In one embodiment, gastric cells, such as the MKN-74 cell line, are seeded on the top channel at a density of l, 0x10⁵ cells/mL and allowed to adhere for 30 min before re - initiating perfusion of cell culture media (Figure 4) .

For mechanical actuation, after confluency of the epithelial monolayer, mechanical actuation is started by application of vacuum on the actuation chamber. Vacuum controlled actuation is managed by the Elveflow system, applied at a frequency of 0,15Hz for a varying period, ranging from 3to 8 days of actuation (Figure 5a-d) .

Device perfusion under operation is conducted at 37 °C, under high humidity and normal atmosphere. Cell culture media perfusion is done at a rate of 2 μL/min (Figure 4) .

Thus, the present invention refers to a method for fabrication of microactuator and in-line air bubble degasser system for organ-on-a-chip devices characterized by comprising the steps of: a) Rendering a device comprising 2 or more sheets designed to create 1 or more chambers; b) Cutting-out the generated sheet designs of laminated PDMS with a thickness of 250 to 500- μm, most preferably 250 μm; c) Mounting each plotted layer on top of the preceding layer and bonding them together through plasma activation of the surface; d) Bonding the assembly to a glass slide; e) Priming the surface, seeding cells and perfusing the fluidic portions of the sheets; f) Degassing air bubbles in-line by applying a constant vacuum pressure to the chamber, most preferably of -50 mbar; g) Cell stretching by vacuum-controlled actuation on the chamber.

In another embodiment, the said seeding of cells is performed by perfusing fibroblasts embedded in collagen type I and an epithelial cell suspension, such as the human derived epithelial gastric cancer cell line MKN74 (which can be purchased from the Japanese Collection of Research Bioresources cell bank) .

In another embodiment, the said seeding of cells is performed by flipping the device over and perfusing endothelial cells, creating an epithelial- endothelial barrier interface.

In one embodiment, the said seeding of cells is performed by bonding perforated PET membranes in between the flow chambers, serving as the cell culture surface, creating cell to cell interactions.

By developing the above-mentioned method, a microactuator and in-line air bubble degasser system for organ-on-a-chip devices can be fabricated. Thus , the present invention also refers to a microactuator and in-line air bubble degasser system for organ-on-a-chip devices characterized by compris ing 2 or more sheets des igned to create 1 or more chambers .

In another embodiment , the said sheets are characteri zed by comprising laminated PDMS with a preferable thickness of 250-500μm .

In another embodiment , the said sheets are characterized by further compris ing a glass slide .

In another embodiment , the said sheets is characterized by also compris ing seeded cells .

In one embodiment , the said seeded cells are characteri zed by comprising fibroblasts , epithelial cells such as the human derived epithelial gastric cancer cell line MKN74 , endothelial cells and combinations thereof .

In another embodiment , the said microactuator and in-line degasser chamber is characterized by compris ing a coupling to vacuum to apply a constant pres sure , most preferably of -50 mbar .

In another embodiment , the said microactuator and in-line degasser chamber is characteri zed by compris ing coupling to vacuum to control mechanical stretching of cells .

In other embodiments , the said microactuator and in-line degasser chamber comprise 2 or more independent chambers .

In conclus ion, the disclosed method generates a system capable of sustaining long term operation under cell culture conditions . Furthermore , it can be adapted to complex organ-on-a-chip designs , by intercalating porous membranes , mechanical actuators and in-line degas sers . The present invention establishes the steps of an innovative method to fabricate complex multi-layered PDMS fluidic devices with integrated microactuators and air bubble degasser that relies in the machining of PDMS laminates using a benchtop cutting plotter . The technique simplifies the entire fabrication procedure. The fabrication process demonstrates the ability to sustain a long-term culture of an epithelial gastric cell line (MKN74 ) , with integration of microactuator within the chip, reproducing biomechanical cues at physiological levels, as well as ability to operate as in-line degassers to eliminate on-chip air bubbles.

Other examples

To characterize the fabrication method, 3 microdevices were designed (Figure 6) . Device 1 is a 5-layered fluidic system, with 2 fluidic linear channels measuring 12 mm in length, arrayed in a cross format and separated by a PET perforated membrane (Figure 6A) . Device 2 is built from 5 superimposed layers. It is composed by 3 alternating layers of 500 μm PDMS laminate foil and a cell culture treated PET membrane (Figure 6B) . Fluidic features are rectangularly shaped, 2.0 mm (width) x 10.0 mm (length) x 0.5 mm (height) , for a total cell culture area of 0.2 cm² and a total volume of 0.01 cm³ (approximately 10 μL ) . Access to the cell culture area is made through a linear feature connecting to an access porthole of 2 mm in diameter (FIG. IB) . Fluid flow is applied unidirectionally, serving one porthole as an inlet and the other as an outlet for liquid ejection. Device 3 was adapted from Device 2, by adding a lower portion, containing the pneumatic actuation system. The actuator is a two-piece assembly comprised of a 0.25 mm(height) flexible PDMS membrane and an actuation chamber with 35 mm (width) x 11 mm (length) ( Figure 6C ) .

Brief Description of the Figures

Figure 1: CAD drawing of each PDMS laminate layer composing the biochip.

Parts a) , b) and c) represent the top, middle and bottom portion of the fluidic assembly respectively, cut from 500 μm PDMS laminate. Parts d) and e) constitute the pneumatic actuator, cut from 250 μm PDMS laminate. All measurements in mm.

Figure 2: Exploded view of the biochip assembly. From top to bottom, each PDMS layer (grey) constituting the fluidic portion of the chip is separated by a perforated PET membrane (pink) . The actuation chamber (*) , is topped by a 250 μm PDMS membrane (yellow) . The entire assembly is mounted on a microscope glass slide (transparent layer) . (Note: The top is a cast-on PDMS layer) .

Figure 3: In-line air bubble degassing. After priming, the chip was filled with PBS solution and allowed to equilibrate at 37 °C for at least 2 h. Air bubbles were injected through the system and pushed until they were inside the cell culture chamber. A constant vacuum pressure of -50mbar was applied on the microactuator and the air bubbles were photographed every hour under a stereomicroscope (Olympus SZX10) coupled with a camera (Olympus EP50) , in order to monitor air bubble area variation. A) In-line air bubble degassing by application of a constant -50 mbar vacuum pressure at the microactuator over a period of time. (B) Area of each bubble was measured using Fiji software and plotted as % of area decrease over-time. Results presented as average measurements ± SD.

Figure 4: Phenotype of MKN74 cells. Metabolic activity was assessed by the resazurin assay. Approximately 1000 cells/well were seeded in a 24- well plate and allowed to adhere for 24 h. Metabolic activity was assessed every 48h for 15 days. For the measurement of metabolic activity, a stock solution of 0.1 mg/mL resazurin (Sigma) was diluted to 20% (v/v; resazurin/cell culture media) . Culture media was replaced with 500 μL of this solution and incubated at 37 °C for 2 h. 200μL of the resulting supernatant were transferred to an opaque 96-well plate with clear bottom (Greiner) and the fluorescence signal was measured (Ex 530 nm/Em 590 nm) in a fluorimeter (model Sinergy MX HM550; Biotek Instruments) . Metabolic activity was expressed as average relative fluorescence units (RFUs) ± standard deviation (SD) . The population doubling time was estimated from the metabolic activity and results registered between 0 and 48 h, at which point cells were non-conf luent and growing at an optimal exponential rate. Immunohistochemistry against Ki67, a marker of proliferation, was performed on a non-conf luent population, 48 h post seeding. Briefly, cells were washed with PBS (3x 5 min) and fixed with 4% (v/v) paraformaldehyde suspended in PBS (20 min at room temperature) . Fixation was followed by further washing with PBS (3x 5 min) , followed by incubation in 50 mM NH₄CI (10 min) and permeabilization with 0.2% (v/v) triton-XlOO (5 min) . Cells were rinsed with PBS ( 3x 5 min) and blocked in 5% (v/v) bovine serum albumin (BSA) in PBS (30 min) . Primary antibody incubation was performed with rabbit anti-Ki67 (dilution 1:100 in 5% (v/v) BSA; Abeam) overnight at 4 °C. Secondary reaction was done with an anti-rabbit alexa fluor antibody (dilution 1:500; Thermo Scientific) for 2 h at room temperature. After rinsing with PBS (3x 5 min) , cellular preparations were mounted in vectashield containing 2- ( 4-qmidinophenyl) -1H-indole444 6-carboxamidine (DAPI) , for nuclei staining. Image acquisition was done in a SP5 confocal microscope (model SP5; Leica Microsystems) .

(A) Comparison study of phenotype of MKN74 cells growing under standard conditions with cells growing over silanized surfaces or under CO2 depleted conditions. (B) Ki67 staining of the same populations. (C) and

(E) show population doubling time. (D) Comparative study of the metabolic activity of cells growing on a non-silanized vs. silanized surface and

(F) cells growing under normal or CO₂-depleted conditions. All graphical representations display average measurement ± SD.

Figure 5: On-chip microactuator. To assess surface expansion, a biochip (Device 3) , was filled with cell culture media and connected to the piezoelectric pressure controller using PEEK tubing. The magnitude of negative pressure applied was managed by the pressure controller, by creating a vacuum generated pull on the flexible PDMS membrane that, in turn, displaced the perforated cell culture membrane above it. By applying mechanical distension from below, the membrane is actuated three dimensionally, by pulling and releasing in the x, y and z planes. Cell culture substrate stretching was assessed at 0, -50, -100, -150 and - 200 mbar. Still images were taken with an inverted microscope (Olympus 1X71) , at rest and actuated state and the average interpore distance (distance between two adjacent pores) , in both states, was measured using image analysis software (Fiji) . Linear expansion was calculated as the amount of linear stretch sustained by the substrate according to the following equation: ε_lin= (L_f-L₀) /L₀, where L₀ and L_f are the interpore length before and after actuation, respectively. Surface expansion was estimated from the linear expansion using the following formula: ε_SA ( ε_lin+1 ) ^2_1 , where ε_SA is the surface expansion and ε_lin is the linear expansion. The results were plotted as percentage of ε_SA ± SD relative to the rest state.

(A) Illustration of on-chip microactuator. (B) Vacuum pressure is applied at the actuation chamber, resulting in 3D cell surface expansion. Colour- coded outlines exemplify cellular expansion for a single cell at different actuation pressures. (C) Surface expansion as estimated from the linear expansion, at different pressure steps. Red line represents average surface tension of in-vivo gut cells. (D) Assessment of microactuator delamination over a period of 10 days. Top plates show a new chip prior to use. Red arrows point to the limits of the microactuator. Part of the flow channel network can be seen out of focus. The bottom plates show the detail of a chip that has been run for 10 days. No delamination of the microactuator is observed (red arrows) . Fluidic channel with seeded MKN74 cells can be seen out of focus, above the microactuator.

Figure 6: Other devices: Exploded 3D view of (A) Device 1, (B) Device 2 and (C) Device 3, each depicted with a top-down view of the assembled structure to the right and corresponding measurements for the main fluidic and microactuator features. Values are in millimeters.

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Lisbon, 12^th August, 2021

Claims

1. Method for fabrication of microactuator and in-line degasser system for organ-on-a-chip devices characterized by comprising the steps of: a) Rendering a device comprising 2 or more sheets designed to create 1 or more chambers; b) Cutting-out the generated sheet designs of laminated PDMS with a thickness of 250 (most preferably) to 500μm; c) Mounting each plotted layer on top of the preceding layer and bonding together through plasma activation of the surface ; d) Bonding the assembly to a glass slide; e) Priming the surface, seeding cells and perfusing the fluidic portions of the sheets; f) Degassing air bubbles in-line by applying a constant vacuum pressure to the chamber, most preferably of -50 mbar; g) Cell stretching by vacuum-controlled mechanical actuation on the chamber.

2. Method according to claim 1, in which the said seeding of cells in step e) is performed by perfusing fibroblasts embedded in collagen type I and an epithelial cell suspension, such as the human derived epithelial gastric cancer cell line MKN74.

3. Method according to claim 1, in which the said seeding of cells in step e) is performed by flipping the device over and perfusing endothelial cells, creating an epithelial - endothelial barrier interface.

4. Method according to claim 1, in which the said seeding of cells in step e) is performed by bonding PET membranes in between the flow chambers, serving as the cell culture surface, creating cell to cell interactions.

5. System for organ-on-a-chip devices characterized by comprising 2 or more sheets designed to create 1 or more chambers.

6. System according to claim 5, in which the said sheets are characterized by, comprising laminated PDMS with a preferable thickness of 250-500μm.

7. System according to claims 5-6, in which the said sheets are characterized by further comprising a glass slide.

8. System according to claims 5-7, in which the said sheets are characterized by further comprising seeded cells.

9. System according to claims 5-8, in which the said seeded cells are characterized by comprising fibroblasts, epithelial cells such as the human derived epithelial gastric cancer cell line MKN74, endothelial cells and combinations thereof.

10. System according to claims 5-9, in which the said in-line degasser chamber is characterized by comprising a coupling to vacuum to apply a constant pressure, most preferably of -50 mbar.

11. System according to claims 5-10, in which the said microactuator and in-line degasser chamber is characterized by comprising coupling to vacuum to control mechanical stretching of cells .

12. System according to claims 5-11, in which the said microactuator and in-line degasser chamber comprise 2 or more independent chambers .

Lisbon, 12^th August, 2021