A Rapid, Reliable and Less-destructive On-chip Mass Measurement for 3D Composite Material Testing Microstructures Gilgueng Hwang 1,2,3* , Christophe David 3 , Alisier Paris 3 , Dominique Decanini 3 , Ayako Mizushima 4 , Yoshio Mita 1,2 1 LIMMS-CNRS, Institute of Industrial Science, University of Tokyo, Japan 2 Dept. of Electrical Engineering and Information Systems, The University of Tokyo, 113-8656 Japan 3 C2N-CNRS, University Paris-Sud, University Paris-Saclay, Palaiseau, 91120 France 4 VLSI Design and Education Center, The University of Tokyo, 2-11-16, Yayoi, Bunkyo-ku, Tokyo 113-0032, Japan * Email: hwang@if.t.u-tokyo.ac.jp, gilgueng.hwang@c2n.upsaclay.fr Abstract-Mass measurements of complex three-dimensional microstructures are essential to the design and development of microsystems. However, the mass measurements of such microstructures require several manipulation steps which are not well controlled. We have demonstrated a rapid, reliable and lessdestructive on-chip mass measurement method by incorporating with AFM micromanipulation. It is based on AFM pick-measureplace micromanipulation using Van der Waals attraction and the mass measurement by resonant frequency shift. The measurement sensitivity revealed to be 25 Hz/pg and it could be promising to characterize MEMS with complex geometries and composite materials.

I. INTRODUCTION

Recent advancement of micro/nanolithography technologies allows to fabricate 3D helical microstructures which are promising building blocks towards micro electromechanical system (MEMS) sensors and actuators. The self-scrolling technology has been studied by strain-engineering of epitaxially grown semiconductor bilayers [1,2]. Various other materials from metals to dielectrics could also be used by glancing angle deposition [3]. Direct laser writing by two-photon polymerization simplified 3D microfabrication [4] and allowed to fabricate 3D gold-helix photonic metamaterial by electrolytic growth [5] or deposition by sputtering [6]. However, such additional processes cause the volumetric variation in the final volume and mass from the initial design thus expected to optimize the process. This requires mass estimation of such composite 3D materials. The computational model estimation is limited to the simple models and could derive with errors from the real value. Therefore, experimental characterizations remain to be pertinent. The major challenge of measuring the mass of fabricated 3D composite microstructures is due to their fragile mechanics which can cause the loss of materials during the measurement.

In this work, we propose a rapid, reliable and less-destructive method to measure the mass of such 3D microstructures. The method is performed by an atomic force microscope (AFM) micromanipulation steps combining both the tapping mode and the contact mode. Using AFM as micromanipulator as well as mass sensor, mass measurements can be done on the fabrication substrate of the target samples. Compared to conventional approaches such as MEMS mass sensor or in-situ scanning electron microscope (SEM) nanomanipulation, this on-chip measurement can avoid additional transportation process from fabrication substrate to MEMS mass sensor (Table 1). The paper describes first on the fabrication of 3D composite material microstructures fabricated on the supporting pillars. Then we describe the chemical-free micromanipulation strategy for the mass measurements with minimal addition or loss. Finally

II. 3D COMPOSITE MATERIAL MICROSTRUCTURES

To demonstrate such characteristics, it is important to prepare samples for this process. We fabricated the triple bihelical test microstructures with polymer and metallic layers which are promising building blocks towards the wireless electromagnetic actuators or sensors [7] (Fig. 1). The helical structures were made by two-photon polymerization using 3D laser nanolithography system (Nanoscribe GmbH). The helical test structures are suspended onto the conical micropillars which elevate 5 m above the substrate surface thus to avoid shadow effect during deposition of other layers. The metallic layers (Cr/Ni 10/200nm) were deposited by sputtering to maximize the deposition uniformity. The uniformly coated layers with minimum non-metalized area (the point contact with conical micropillars) and the conical supporting pillars from the

Short (direct measurement)

Table 1: Comparison of the proposed method with others substrate allow to detach easily the helical structure by AFM cantilever (Fig. 1) without losing the mass.

III. MICROMANIPULATION OF TESTING MICROSTRUCTURE

Since the fabricated microstructures are fixed on the substrate, several micromanipulation steps are required to prepare for the mass measurements. This includes the detachement out of substrate and the attachment to AFM tip.

A. Detachement of Microstructures from Substrate

To demonstrate such characteristics, it is important to prepare samples for this process. We fabricated the triple bihelical test microstructures with polymer and metallic layers which are promising building blocks towards the wireless electromagnetic actuators or sensors [7] (Fig. 1). The helical structures were made by two-photon polymerization using 3D laser nanolithography system (Nanoscribe GmbH). The helical test structures are suspended onto the conical micropillars which elevate 5 m above the substrate surface thus to avoid shadow effect during deposition of other layers. The metallic layers (Cr/Ni 10/200nm) were deposited by sputtering to maximize the deposition uniformity. The uniformly coated layers with minimum non-metalized area (the point contact with conical micropillars) and the conical supporting pillars from the substrate allow to detach easily the helical structure by AFM cantilever (Fig. 1) without losing the mass.

The detachments of the microstructures out of the fabrication substrates are important but conventional micromanipulation approaches are not free from the addition or loss of mass mainly due to the requirement of brutal manipulation or chemical solutions. Therefore, here we propose a chemical-free and controlled detachment process. The specially designed 6 supporting pillars with minimum surface contact area with testing microstructures. The pillars are conical to keep mechanical stability to hold the testing microstructure during fabrication and transport processes. And it also helps to minimize the contact area which is the main source of potential mass addition or loss during the detachment process.

In addition, we attempted to control better the micromanipulation process to detach the testing microstructure. AFM tip is engaged near to the supporting structure to avoid any potential damage. Then the tapping mode with scanning the interface area to the supporting structure. This is to be sure to precisely place the AFM probe tip on top of the bottom supporting ring. Without having this process step, the detachment can include the supporting structure which results in a wrong measurement. Figure 3 shows the scanned image and photograph of the AFM tip adjacent to the supporting ring. Based on the AFM topography, the AFM tip is moved on to the supporting ring. Then the supporting pillars are disconnected by moving further towards the center of the ring.

The disconnected microstructure is attracted electrostatically to the AFM tip and attached to it. Once it is attached, the AFM tip is moved vertically to Z axis to free the microstructure out of the substrate. We perform the mass measurement at this step. Once the measurement is done, we place the microstructure to the target substrate. Figure 3 shows the sequence of manipulation.

B. Attachement of AFM Probe and Microstructure

It is also crucial to avoid any possible addition or loss of mass during the mass measurement process. One of the major sources of the addition of mass could be caused from the addition of chemical glue dispensed to fix the AFM tip and the testing microstructure. For this concern, we chose to use Van der Waals force for the attachment. This will allow the attachment to be reversible. In addition, electrostatic force induced from the charged AFM tip helps to hold the attachment firmly during the handling and the measurement.

Once the measurement is completed, the tested microstructure can be released out of the AFM tip for the next measurements. This can be made by placing the microstructure onto the substrate thanks to the fact that the microstructure`s contact surface area to the substrate is higher than the one to the AFM tip (Figure 4). Moreover, the engagement of the AFM tip with the substate will apply a minimal force sufficient to deflect the helical microstructures attached to the AFM tip thus allows to break the contact between the AFM tip and the microstructure. The remaining surface charge induced electrostatic force can also be reduced during the surface contact mode which makes easier the release process of the microstructures.

IV. MASS MEASUREMENTS OF MICROSTRUCTURES

The micromanipulation allows to handle testing microstructures. In this section, mass measurements and further analyses of the testing microstructures are described.

Prior and posterior to the mass measurements, we will need to perform following micromanipulation. To briefly summarize, the calibrated AFM tip is approached near to the bottom supporting part of the target sample. Surface scanning by AFM tip gives topographical information at the interface between the support ring and substrate. This allows to optimize the precise positioning of the AFM tip for the further releasing process (Fig. 2). Then a sequence of pick-measure-place is performed by contact mode micromanipulation of AFM tip (Figs. 1 and2). The AFM switches back to tapping mode after the target sample is picked up in air and another thermal calibration is performed to obtain resonant frequency of the AFM cantilever with added mass of target sample. The addition of a small mass ǻm of compared to that of the cantilever causes the shift in the resonant frequency ǻf. With the constant stiffness of the cantilever, the resonant frequency variation is given by the first order approximation [8]:

ǻf = -2ʌ 2 •ǻm•f 3 /keff ( 1 )
Once the measurement is made, the AFM tip with the 3D helix moves back to the substrate to place the measured samples back for other measurements (Fig. 4). This pick-measure-place can be repeatedly performed with same sample or different samples and resonant frequencies are measured at each time (Fig. 5). The 6 different types of samples with slight geometrical modifications were tested to reveal the mass of each and the measurement sensitivity (Table 2). The mass was ranging from 1334 pg to 2142 pg and the sensitivity was revealed to be 25 Hz/pg. Iterative measurements of identical testing microstructure with pick-and-place manipulations showed that the deviation of the measured mass is within the sensitivity.

Further analyses of the deposited layer thickness were also able to be made and it could be useful to identify the measured mass of inner resist layer and the deposited layers (Figure 6). AFM tip scanned on an area over the interface between the masked (not deposited) area and the deposited area. This reveals also the deposition quality by surface roughness. The deposited metal layer (aimed to deposit Cr/Ni 10/200nm) thickness was measured around 216 nm (difference is less than 3%). The roughness of the surface was measured around 2 nm (less than 1% of the total thickness). These are within the acceptable range of the quality of the additionally deposited layers.

Moreover, defects or morphological inspections can also be performed by rolling the 3D microhelix (Figure 7). For this purpose, we placed testing structure laterally on the substrate and used AFM tip to push to roll the structures. The analyses by SEM photographs reveal that the testing microstructures did not have any major deformations or damage during the whole manipulation and measurement processes. This makes the proposed process to be promising to the mass analyses of 3D complex and composite microstructures.

V. CONCLUSIONS

We proposed a mass measurement method by AFM for 3D complex shape composite microstructures. Specially designed supporting micropillars allowed the microstructures can be easily detached from fabrication substrate and attached to AFM tip for the mass measurement by resonant frequency shift. Thanks to the purely physical (chemical-free) micromanipulation, the measurement process can be made iteratively. The tested microstructures can be conserved for more measurements and also for further analyses such as electron microscopy. The proposed on-chip mass measurement method is shown to be rapid, reliable and less-destructive thus could be promising to test MEMS with 2D or 3D complex geometries and composite materials.

Figure 1 :

 1 Figure 1: 3D helical test microstructure and AFM cantilever. Schematic concept of pick-measure-place testing structures.

Figure 2 :

 2 Figure 2: Scanning topography around the bottom supporting ring for a fine position.

Figure 3 :

 3 Figure 3: Pick-measure-place micromanipulation by AFM: 1. AFM tip is placed on top of the bottom support ring between two supporting pillars and move towards the center of microhelix, 2. AFM tip is moved laterally to disconnect the microhelix from supporting pillars, 3. Retraction of AFM tip out of the microhelix, 4. Moving up the picked-up microhelix with AFM tip, 5. Verification of the remained bottom supporting ring, 6. The microhelix is placed on to other substrates for further analyses.

Figure 4 :

 4 Figure 4: Placing back the 3D microhelix onto substrate.

Figure 5 :

 5 Figure 5: Mass sensing by resonant frequency shift

ACKNOWLEDGMENT A part of this work is supported by Japan Society for the Promotion of Science (JSPS) through invitational fellowships for research in Japan. The test microstructure design, 3D nanolithography and AFM measurements were supported by the French RENATECH network through its C2N technological facility. Metallic layer deposition by sputtering was supported by MEXT/VDEC Nanotechnology Platform in The University of Tokyo.


Alisier Paris

Dominique Decanini

Gilgueng Hwang
email: gilgueng.hwang@c2n.upsaclay.fr

Swimming force characterizations of multistaged bi-helical microswimmer and 3D vortex trap manipulation

Keywords: Helical microswimmer, multistaged bi-helix, two-photon 3D lithography, magnetic actuation, noncontact 3D manipulation, hydrodynamic vortex trap

Swimming force characterizations of multistaged bi-helical microswimmer and 3D vortex trap manipulation

Alisier Paris, Dominique Decanini, Gilgueng

MAIN TEXT

Introduction

The manipulation of microorganisms (bacteria, yeasts, etc.) or individual cells has always been a challenge and a limitation for biologists and physicians. Yet it has been shown that mechanical interactions play an important role in cell life [START_REF] Lintilhac | Stress-induced alignment of division plane in plant tissues grown in vitro[END_REF]- [START_REF] Helmlinger | Solid stress inhibits the growth of multicellular tumor spheroids[END_REF] . Thus, having an effective manipulation tool should open up many opportunities that go far beyond simple study and organization. For example: the measurement of the stiffness of a cell wall can predict the risk of cancerous transmission [START_REF] Swaminathan | Mechanical Stiffness Grades Metastatic Potential in Patient Tumor Cells and in Cancer Cell Lines[END_REF] and the application of mechanical force to a tumor can control its evolution [START_REF] Itabashi | Mechanical impulses can control metaphase progression in a mammalian cell[END_REF] , the modulation of the force applied on a cell makes it possible to play on its mitosis [START_REF] Itabashi | Responses of chromosome segregation machinery to mechanical perturbations[END_REF], [START_REF] Knouse | Chromosome Segregation Fidelity in Epithelia Requires Tissue Architecture[END_REF] and thus on the mechanisms of genetic reproduction, or, more simply, the manipulation of gametes can greatly facilitate in-vitro reproduction. All these examples require having a very precise manipulation tools that can act in three dimensions at the microscopic scale in a highly controlled environment. Also, over time, different techniques have emerged to allow such operations. The main and most known being atomic force microscopy (commonly called AFM), optical tweezers and magnetic tweezers [START_REF] Neuman | Single-molecule force spectroscopy: optical tweezers, magnetic tweezers and atomic force microscopy[END_REF] . However, these methods are generally invasive or destructive for the biological material studied, which limits or prohibits their use in many cases. AFM was initially conceptualized to operate in a vacuum and nanoscale. Although remaining usable in the air it requires heavy adaptations to work in a liquid and the working scale is rarely adapted to cell biology work. In addition, the piezoelectric base tip handling infrastructure is too large to fit into a microfluidic system. On the other hand, optical tweezers seem ideal to be combined with microfluidic systems, the latter being generally made of transparent materials. However, the use of high intensity focused electromagnetic fields tends to locally heat the medium and damage the manipulated object [START_REF] Neuman | Characterization of Photodamage to Escherichia coli in Optical Traps[END_REF] . In the case of work on fragile biological material, this damage can be significant. Finally, magnetic tweezers can not, with rare exceptions [START_REF] Blakemore | Magnetotactic bacteria[END_REF] , be used with biological material without attaching to or in this material magnetic field sensitive bodies. It is quite obvious that this makes it a very intrusive technique, which is often harmful for studies in biology.

The wide variety of possible actions with a micro-robot -manipulation and movement of material in 2D and 3D by direct mechanical forces application (by pressing or pushing) [START_REF] Kawahara | On-chip microrobot for investigating the response of aquatic microorganisms to mechanical stimulation[END_REF]- [START_REF] Zhang | Reliable Grasping of Three-Dimensional Untethered Mobile Magnetic Microgripper for Autonomous Pick-and-Place[END_REF] or indirect ones (via hydrodynamic flow generation) [START_REF] Floyd | Two-Dimensional Contact and Noncontact Micromanipulation in Liquid Using an Untethered Mobile Magnetic Microrobot[END_REF]- [START_REF] Tung | Non-contact manipulation for automated protein crystal harvesting using a rolling microrobot[END_REF] -make it a very versatile tool having a wide range of possible applications. The microscopic swimmer moving on a smaller similar scale biological sample, and study material should be able to apply force at scales similar to the stresses naturally supported by this material, and thus avoid any risk of damage. Especially since microrobots need only a weak magnetic field to function -less than ten milliTesla. And it integrates naturally into a microfluidic device, offering a very suitable environment for work in biology.

Our objective is to develop a new non-contact micromanipulation tool towards applications in cells studies. It should allow precise manipulation and characterization of cells as much individual than at group, and the possibility to have automated functions. As first step we've proposed a design for a rotating helical microswimmer capable of generating vortical flow which allows highly stable noncontact manipulation possibilities [START_REF] Paris | On-chip multimodal vortex trap micro-manipulator with multistage bi-helical micro-swimmer[END_REF] . But this version lost the ability to move, and manipulate, in 3D. In this paper we propose the improvement applied to this swimmer to recover the 3D mobility, and so allowing non-contact 3D manipulation of particles. As for previous work we use microfluidic chip as controlled and stable workspace for the development of this tool. It will be our in-vitro environment, and will be useful for future use with biological material.

In this paper we will explain first how we have evolved design and fabrication process of the swimmer. Then we will develop characterization of 6 different models of swimmer we have conceptualized. And finally we will show 3D handling and manipulation promising toward future biological applications.

Materials and methods

Design and modeling

The major idea of our previous works [START_REF] Paris | On-chip multimodal vortex trap micro-manipulator with multistage bi-helical micro-swimmer[END_REF] is to inverse the central part of the helical micro-swimmer.

We expected that this inverted part made local counter-flux opposed to the global hydrodynamics around the robot. This generates local vortices near the robot which enable to trap particle. As the first tests show that the vortex effect is too small compared to the propulsion power of these models.

We chose to add a functional head at each side of the models, helping stabilization of side trap and allowing head trap. The results of these first experiments confirmed the possibilities of the concept.

However the demonstrated mobility and non-contact manipulation performances are promising but still limited to 2D manipulation. In this section, we show how to overcome this limitation and extended the concept to 3D manipulations.

Table 1. Design evolution

Simple helix [START_REF] Barbot | On-Chip Untethered helical Microrobot for Force Sensing Applications[END_REF] RT3 [START_REF] Paris | On-chip multimodal vortex trap micro-manipulator with multistage bi-helical micro-swimmer[END_REF] RT4 (this work)

Refined model

The proposed design in this work was aimed to enable 3D propulsion and manipulation while keeping the major functionalities for the 2D non-contact manipulation which are multimodal (including both the head and side entrapment), stable and non-contact manipulation shown in the previous work [START_REF] Paris | On-chip multimodal vortex trap micro-manipulator with multistage bi-helical micro-swimmer[END_REF]. In addition, we kept the bi-helix design meant for the double helical geometry inspired by DNA`s double helix. The bi-helix contributes to improve the propulsion stability and smooth response of the swimmer to the magnetic field by reduced wobbling effects due to the deviation of magnetization axis to the axis of rotation especially at low propulsion frequency regime.

More details could be found on our previous work [START_REF] Paris | On-chip multimodal vortex trap micro-manipulator with multistage bi-helical micro-swimmer[END_REF].

Figure 1: Schematic concept of the proposed system with the evolved design and fabrication from RT3 to RT4. The proposed concept is to enable 3D propulsion while keeping multimodal noncontact manipulation capabilities. The overall dimensions of the RT4 model take the dimensions of the "D" variants of the RT3 model [START_REF] Paris | On-chip multimodal vortex trap micro-manipulator with multistage bi-helical micro-swimmer[END_REF]. Otherwise, we made three main changes in the design. First of all a global refinement of the structure to lighten the swimmer. Secondly we do not find on the RT3 that complex design for the head have a real utility, so we reduce it to its simplest form : a ring. Finally we found that breaks in hydrodynamic generated by the heads is too strong in RT3 due to brutal transition so we design the external propeller to make a slope between the head and the central propeller.

Figure 1 shows the design modification to recover 3D propulsion and manipulation. Starting from the RT3 variant which had the best results we took the general geometry ratings (number of helix, number and length of pitch) for the realization of the next model, named RT4. In order to recover the 3D displacement, we modified some geometrical parameters by relying on several hypotheses: The first, reasonable one, is to think that the simplest approach to improve the propulsion / weight ratio remains to refine the swimmer, and therefore reduce its weight. In that case the lower limit of resolution is the voxel size of 3D nanolithography by two-photon absorption (Nanoscribe, GmbH).

Also, the 3D model has been reduced to a minimum, and the bulk of the swimmer's volume comes from the voxel via the product of convolution of the latter with the 3D model. This results in propellers with a thickness of 0.2 μm. In addition, Since the utility of a complex head as for the RT3 model has not really been proven, we have decided to make a simple ring with the thickness of 0.5 μm, outer diameter 16 μm and width 3 µm. A second hypothesis assumes that the brutal cuts in the RT3 model, especially at the head, can only perturb the generated vortical flow, and thus weaken the net propulsion. We therefore smoothed the swimmer's geometry so that the outer propellers made a slope between the head and the central propeller, as shown in Figure 2. Variants with "LL" nomenclature are variants with wide propellers where the variants with "LS" nomenclature are those with fine propellers. This distinction is illustrated with red ribs. The next three digits represent the outer radius in micrometer of each piece of the helix, so "838" means that the helix has a radius of 8 μm at the heads and 3 μm at the central helix . Likewise "848" implies that the inverted central portion has a radius of 4 μm while keeping the lateral ends at 8 μm. Finally the "A" implies the presence of a reinforcement ring at the junction between the propellers, as shown in green on the right. There are two variants not illustrated here, the LS_838 and LL_838, which are exactly the same as, respectively, the variants LS_A838 and LL_A838 with the exception of the absence of the reinforcement ring.

However we have comparison as main source of discrimination, which requires creating several variants around the model. A first variation allows us to study the effect of the width of the propellers, the parameter for the moment neglected. We have therefore separated the variants into two categories: a category LS having helices of a width equivalent to the ring forming the head, i.e. 3 μm, and another LL having double-width helices, namely 6 μm. A second comes from a concern about the strength of the attachment between the lateral and central propellers. We are interested here in the difference between a variant "naked" and another having a reinforcement ring (variant "A"). A third is offered by the presence of the reinforcement ring, which allows us to enlarge the outer diameter of the central helix from 6 μm (variant "838") to 8 μm (variant "848"). These variations are shown in Figure 2. On the left, the method used for the RT3 model. The resin core being printed "lying down" on the substrate (in indigo), the rotation (in orange) of this support according to the normal to its surface (in green) is not sufficient to suppress the shadowing effect during the deposition metallic (in red). In the middle, a view of the spirit of the swimmer's area covered by the metal deposit is indicated in red. In the middle right, according to the same color code as on the left, the situation with the model RT4 printed "standing". On the right, the SEM image of the LL_A838 variant printed vertically and metallized by sputtering.

Vertically standing structures

A consequent defect requiring a lot of anticipations for mixed results coming from the 3D lithography process. The crosslinking of the resist is made on a non-spherical voxel in the form of a prolate spheroid of 1.0 µm over 0.4 µm. The volume of this voxel being considered as an infinitesimal point by the printing software, it generates strong deformations when the details of the model are near the micrometer as the final morphology is the product of convolution between the 3D model and this voxel. In the framework of the RT3 model, we anticipated this distortion on the 3D model. This deformation is essentially harmful in the context of an impression of the swimmer "lying down" because it is not uniform by rotation around the geometrical axis of the helix. So we decided to pick up the swimmer and print it vertically. This decision required the modeling of a new medium but allowed two major advantages. The first is a considerable reduction of the harmful effect of the deformation since it becomes homogeneous in rotation. This allows us to no longer have to torture the 3D model before the printing phase and ensure better reproducibility of this model in the resin.

The second, which is related to the next point, is that this configuration no longer masks a swimmer's side to coating processes.

Uniform coating

Beyond the deformation of the swimmer's resist core, it is important to consider the metallization processes, essentially for the ferromagnetic layer. The previous models saw their ferromagnetic layer made by physical vapor deposition by electron beam evaporation (EBPVD). This technique deposits the material at a fixed angle, which, in the case of our swimmers, makes the ferromagnetic layer grow on one side of the resist core. This results in offsets between the centroid of the total volume and the volume of ferromagnetic and the center of inertia of the swimmer, and therefore instabilities. In an attempt to reduce this effect, the deposition was made with a rotating angle, however the swimmer's extended position produces shadowing phenomena that are difficult to compensate, as shown in Figure 3. The rise of the swimmer detailed previously allowed us to eliminate this shadowing and allow, always with a rotating angle, to achieve a much more homogeneous layer and especially centered on the same geometrical axis as the resin core.

Fabrication process

The swimmers were modeled with a CAD design software (SolidWorks©), then exported to the proprietary format of the 3D nanolithography tool (Photonic professional GT, Nanoscribe GmbH).

The core of the robot is made of epoxy-based polymer resist (IPG 780) by two-photon 3D laser lithography on a glass substrate. To ensure uniform metallization, we decided to have metallization done by cathodic sputtering of two metal layers (Cr/Ni 20/200 nm). This physical vapor deposition technique allows a much more uniform recovery than the EBPVD used previously [START_REF] Mattox | Physical vapor deposition (PVD) processes[END_REF] . Fully exploiting the new vertical geometry of manufacture, this technique theoretically allows us to have a greater homogeneity of the metal layer than for the previous models. Finally, to reinforce the effect of the magnetic field we use the hysteresis of the nickel to impose a permanent magnetization on the swimmer. For this, we applied to the swimmer a magnetic field of 200mT aligned with the geometric axis of the helix.

Our microfluidic environment is reduced to its simplest form, a unique isolated chamber of a cylinder of 1 mm height and 6 mm diameter filled with isopropanol. It is made of a unique layer of PDMS bound on a glass substrate. To allow fluidic filling we made a single little hole of ~100 µm diameter on the side of the chip. Then we integrate manually the swimmer before the plasma bonding of the PDMS on the glass.

Characterizations

As this work is an improvement of our previous one, we will focus on propulsion force as a limiting factor for 3D propulsion and manipulation. In this section, we will show the propulsion characterizations of the microswimmers to reveal the generated propulsion force. And, as we only have indirect information from our experimentation, we have to recover force from gravity compensation, so we have to discuss a mass estimation of our swimmers. With the data of volume and surface extracted from 3D modeling and the knowledge of swimmers manufacturing processes it becomes possible to calculate a theoretical value of the volume and the mass (Table 2) of the swimmers. Despite this we have made some simplifications because some data are inaccessible :

Mass estimation

• The volume of resin partially takes into account the deformation due to the printing voxel according to the formula: Vres=Vmod+Smod×voxel error. We have chosen to consider a uniform voxel error of 0.4 μm although the voxel is oblong because the model essentially consists of a standard surface close to the z axis, the error along this axis has the greatest impact. We are conscious of maximizing the error here, which generates a positive error on the calculated final mass.

• Chromium and nickel volumes were calculated using the same formula: Vmet=Smod×thickness of metal. We are aware here of neglecting the surface increase due to the deformation of the voxel, and in the case of nickel that due to the deposition of chromium, which generates a negative error.

• The resin mass was calculated by producing the product of the resist volume and the density of the crosslinked resist as determined experimentally by Antoine Barbot [START_REF] Barbot | On-Chip Untethered helical Microrobot for Force Sensing Applications[END_REF] :ρresist=1.1028kg•dm -3 .

• Metal masses were calculated by considering their bulk density values.

3D Propulsion

The propulsion of the swimmer is done by a rotating magnetic field generated with Helmholtz coils assembly. Our setup generates a magnetic field of 6 mT. The generated torque on a magnetized body in a magnetic field is given by :

(1)

With µ0 the permeability of free space, the V volume of the body (in our case the volume of the ferromagnetic Ni coating), the magnetization of the body and the applied magnetic field. By rotating the field, we made the swimmer rotating along its helical axis, generating a propulsion flux. This propulsion was studied by Edward M. Purcell [START_REF] Purcell | The efficiency of propulsion by a rotating flagellum[END_REF] and follow the matrix :

(2)

With F the sum of external forces applied to the swimmer, τ the couple applied to the swimmer, v the velocity of the swimmer along his helical axis, ω the rotation speed of the swimmer and A, B and C configuration-dependent constants. Here the external forces are the resultant gravity forces Pv(θ)

(weight and buoyancy) applied to the direction of the velocity vector (with the angle between the weight vector and the velocity vector) and a drag force T=T0 v of its displacement in the fluid. And the torque τ is the torque τm generated by the rotating magnetic field. This changes the equation 2 to:

(3) (4)

Figure 4: Swimmer's 3D characterization. On the inset figure left, a schematic of the manipulation with h the height of the characterization chamber, θ the angle between the normal to the surface and the speed direction, l the achieved trajectory, ν the speed of the swimmer, Pv the resultant gravity force and T the drag force. The measurement time sped to travelled distance l at a given frequency gives the speed v of the swimmer. On the right, the speeds according to the imposed rotational frequency for the six variants of RT4. We can observe that we are seeing a linear form as expected with the equation 5. However the results of variant LS_A848 differ greatly from others and are possibly a measuring artifact. The inset right image shows the 3D Helmholtz electromagnetic coils to generate 3D rotating magnetic field [START_REF] Barbot | On-Chip Untethered helical Microrobot for Force Sensing Applications[END_REF][START_REF] Paris | On-chip multimodal vortex trap micro-manipulator with multistage bi-helical micro-swimmer[END_REF][START_REF] Barbot | On-chip Microfluidic Multimodal Swimmer toward 3D Navigation[END_REF].

If we are in the situation of the couple τm is strong enough to impose its rotation speed to the swimmer, we can consider that ω is a constraint, leave beside the equation 4 and reformulate the equation 3:

(5) Thus we have a linear relation between the swimmer's speed and its rotation frequency as long as we stay in a frequency range where the torque generated by the magnetic field on the swimmer is sufficient to force a rotation frequency equal to that of the magnetic field. In order to determine the parameters C1 and C2(θ) we measured the time that the swimmer has to travel the height of the characterization chamber which is 1mm with an angle θ =12°at different frequencies of rotation. As the measurement was chronometered with the eye, it was necessary to have an angle θ > 0 so that when the swimmer comes into contact with the upper surface of the chip, he changes his mode of movement to the surface motion (spintop) [START_REF] Barbot | On-chip Microfluidic Multimodal Swimmer toward 3D Navigation[END_REF] . This displacement from the mode shift is very visible thus easy to chronometer the travel time. It remains reasonable to consider a measurement error of the order of one second. From these times and a calculation of the travelled distance we can go back to a value of speed dependent on the frequency. These experimental results are summarized in figure 4. A linear regression of these results allows us to find the parameters and presented in Table 3 with a very high degree of certainty, as evidenced by the value of the coefficient of determination R².

Moreover, the knowledge of the mass of the swimmers, studied previously, allows us also to extract the value of the constants B and A-To. These last two constants are essential because they allow us to free ourselves from the dependency of characterization. It becomes possible to calculate the swimmer's 3D swim evolution from these constants and its mass, an essential step for automating 3D mobility. In order to have a basis of comparison in terms of the force generated by the swimmer in rotation we will consider the equation 5 considering the limit case where the swimmer is at equilibrium and therefore at a speed of zero, we obtain the relationship:

Table 3. Constants of equation 5 for an angle of inclination of 12°

(6) 3. There is a clear dominance of the variants LS_838, LL_838 and LL_A838. Once again results for the variant LS_A848 do not have the same evolution as the others which tends to confirm the hypothesis of measurement artifact.

And since we are in the case where the swimmer is in equilibrium, we can consider that in this case Pv(θ) is the force generated by the swimmer. Obviously, experimental results show that the swimmer is not at equilibrium at theses frequencies, which means that is more a minimal propulsion force. But we can therefore reasonably extrapolate F=Bω as a force value generated by the swimmer for comparison between variants. The graph in Figure 5 compares the variants according to the swimmer's force and speed torque running at frequencies of 100, 150 and 200Hz. We confirm that the Figure 5 characterizations were all below the step-out frequencies of the microswimmers. The Figure 4 shows that swimmers have linear reaction to the frequency on the 70-190Hz range while their step-out frequencies are around 250-300Hz for these models. These first results seem to disqualify the variants LS_A838 and LL_A848 which have the worst results. The evolution with the frequency of the results of model LS_A848 seems to go in the direction of the hypothesis of an artifact.

The last three models (LL_A838, LL_838 and LS_838) become difficult to discriminate and offer three options for optimization: the LS_838 will be preferred on applications requiring speed, the LL_A838 on those requiring power and the LL_838 in those requiring a balance between the two.

Simple material manipulation

As for the previous model, we carried out manipulation experiments in isopropanol with 10 µm polystyrene beads. As it is difficult to automate the 3D manipulation, the results were all obtained by manual manipulation of the swimmer.

2D manipulation

The RT4 model seeking to keep the previous model's 2D features, we expected to have the same handling options. The performed tests have made it possible to qualitatively reproduce the three manipulation options detailed in our previous work [START_REF] Paris | On-chip multimodal vortex trap micro-manipulator with multistage bi-helical micro-swimmer[END_REF] . However, if head capture and push / pull options are substantially the same, lateral capture has become more complex. This difference was predictable since we remain on a paradigm of opposition between propulsion flow and counter-flow of capture and that this model sought to improve the propulsion to have a displacement 3D. It has therefore become necessary to increase the rotational frequency of the swimmer to effectively capture the same category of beads. The blue arrows represent the movement of the swimmer. The blue rings highlight the swimmer and the red rings highlight the trapped particle. The black line on the screenshots is the border of a glass plate of 400 µm tickness. (1) Capturing the particle, this can be done in "Spintop" or "Swim" mode depending on the situation. It was done in "Spintop" on the screenshot. (2) Without changing the direction of rotation of the swimmer it is turned 180° on its Pitch angle to switch to "Swim" mode.

(3) The swimmer takes off in "Swim" mode with the captured particle at his head. (4) The swimmer is controlled in "Swim" mode, with all the maneuverability offered by this mode of displacement. ( 5)

The swimmer goes to the place where the particle must be released. ( 6) Once the particle is positioned, the direction of rotation of the swimmer is reversed so that it takes off in the other direction leaving the particle in place.

3D manipulation

The RT3 model had three capture zones, one at the counter-helix and one at each end. We could see in the context of the 2D movement that these capture areas still exist but for the passage in 3D it gets complicated. As explained above, the lateral capture zone is in the state the result of a balance between general flow, counter flow and effect of weight. This works in "Spintop", but reversing the direction of rotation of the swimmer, and thus the direction of flow and counter-flow, to enter "Swim" mode breaks this balance and prohibits any stable capture at this point. Likewise, the area at the rear head -i.e. on the side of the propulsion -can no longer be a stable trapping area since the propulsion of the swimmer will expel the particle. There remains only one usable area in 3D: the head zone is the opposite of the propulsion. It becomes complicated to use this swimmer and the maneuver to be done to capture a particle, shown schematically in Figure 6, can be tedious. However, once the particle is captured and the swimmer in the "Swim" situation, this capture is very stable and allows a totally free 3D manipulation, and a very precise positioning.

In the Figure 6, we can assure that the microswimmer traps a micro bead, moves over the additionally put glass substrate (the height is 400 microns) on the left part (colored part) of the screenshots and successfully placed the bead on top of the glass substrate 400-micron higher than the base substrate on the right side of the screenshots.

Conclusion

We propose a new helical microswimmer model (RT4) for non-contact 3D manipulation. The previous model (RT3) achieved goal of highly stable hydrodynamics capture by generated vortical flow but lost the ability to move in three dimensions, which de facto limited the manipulation options to two dimensions only. The RT4 has had a refinement of its structure and especially a reflection on its manufacturing protocol. Enhancing both the propulsion force by new design and the surface homogeneity of metallic layer coating allowed the RT4 to recover mobility in three dimensions, the "Swim" mode. However, the complexity of the balancing between the propulsion and capture flows for the lateral zone does not make it possible to use it in "Swim" mode, but the capture zone at the level of the swimmer's head allows a stable capture, and therefore a manipulation, in three dimensions.

We keep the side areas for two-dimensional manipulation, with the "Spintop" mode. In this way, this model achieves non-contact manipulation in two and three dimensions by hydrodynamic vortices.

We believed that this will be promising to serve as on-chip micromanipulator with force characterizations of biological samples. To apply the proposed microswimmer for live cell manipulation, we need to carefully demonstrate and characterize the propulsion and manipulation performance in biological culture medium. The remaining question for biological cell manipulation could be the effect of the stretching due to the vortex on the cell may be too strong, but it'll need proper experimentations for the verification.

Figure 2 :

 2 Figure 2: Variants on the RT4 model. The overall dimensions of the RT4 model take the dimensions of the "D" variants of the RT3 model and are shown with black ribs.Variants with "LL" nomenclature are variants with wide propellers where the variants with "LS" nomenclature are those with fine propellers. This distinction is illustrated with red ribs. The next three digits represent the outer radius in micrometer of each piece of the helix, so "838" means that the helix has a radius of 8 μm at the heads and 3 μm at the central helix . Likewise "848" implies that the inverted central portion has a radius of 4 μm while keeping the lateral ends at 8 μm. Finally the "A" implies the presence of a reinforcement ring at the junction between the propellers, as shown in green on the right. There are two variants not illustrated here, the LS_838 and LL_838, which are exactly the same as, respectively, the variants LS_A838 and LL_A838 with the exception of the absence of the reinforcement ring.

Figure 3 :

 3 Figure3: Schematics of the shading problem during metallic deposits and SEM image. On the left, the method used for the RT3 model. The resin core being printed "lying down" on the substrate (in indigo), the rotation (in orange) of this support according to the normal to its surface (in green) is not sufficient to suppress the shadowing effect during the deposition metallic (in red). In the middle, a view of the spirit of the swimmer's area covered by the metal deposit is indicated in red. In the middle right, according to the same color code as on the left, the situation with the model RT4 printed "standing". On the right, the SEM image of the LL_A838 variant printed vertically and metallized by sputtering.

Figure 5 :

 5 Figure 5: Comparative graphs of the RT4 variants according to the Force vs. Speed at given

Figure 6 :

 6 Figure 6: Diagram and screenshots of the steps of capture and positioning of a particle in 3D.The blue arrows represent the movement of the swimmer. The blue rings highlight the swimmer and the red rings highlight the trapped particle. The black line on the screenshots is the border of a glass plate of 400 µm tickness. (1) Capturing the particle, this can be done in "Spintop" or "Swim" mode depending on the situation. It was done in "Spintop" on the screenshot. (2) Without changing the direction of rotation of the swimmer it is turned 180° on its Pitch angle to switch to "Swim" mode.(3) The swimmer takes off in "Swim" mode with the captured particle at his head. (4) The swimmer is controlled in "Swim" mode, with all the maneuverability offered by this mode of displacement.[START_REF] Swaminathan | Mechanical Stiffness Grades Metastatic Potential in Patient Tumor Cells and in Cancer Cell Lines[END_REF] The swimmer goes to the place where the particle must be released. (6) Once the particle is positioned, the direction of rotation of the swimmer is reversed so that it takes off in the other direction leaving the particle in place.

Table 2 . Theoretical volume and mass of swimmers

 2

		Vmodel [µm 3 ]	Smodel [µm 2 ]	Vtotal [µm 3 ]	Mtotale [ng]
	LS_838	298	2034	949	2.88
	LS_A838	318	2123	997	3.01
	LS_A848	329	2351	1081	3.31
	LL_838	374	2444	1156	3.48
	LL_A838	395	2534	1206	3.61
	LL_A848	422	2859	1337	4.05

Acknowledgments

We appreciate Ms. Ayako Mizushima and Prof. Yoshio Mita for their supports to sputtering metallic layers at VDEC, The University of Tokyo during the closed period of C2N cleanroom facility.

Funding

This work was supported by Centre National de la Recherche Scientifque and Universite Paris-Saclay. This work is also partly supported by C'Nano IdF DIM Nano-K and a public grant overseen by the French National Research Agency (ANR) as part of the "Investissements d'Avenir" program (Labex NanoSaclay, reference: ANR10-LABX-0035).


D Decanini

A Harouri

Y Mita

B J Kim

G Hwang

Review of Scientific Instruments 3D micro fractal pipettes for capillary based robotic liquid handling

3D micro fractal pipettes for capillary based robotic liquid handling

D Decanini, A Harouri, Y Mita, B J Kim, Gilgueng

Miniaturized and mobile liquid handling devices are essential elements to biological or clinical applications. This will innovate the conventional liquid handling methods like manual or automated pipetting systems. Here we propose the micro fractal pipette as the candidate device for this objective. It is made of Epoxy polymer and printed by innovative 3D nanoprinting technology based on two-photon absorption polymerization with sub-microns resolution. We demonstrated the efficient liquid handling performance by the micro fractal pipette between the source droplet and the target hydrogel substrate. Thanks to the high porosity up to 78% of total volume and 8.5 times higher cavity surface area compared to the full pyramid with same dimensions. The biomimetic inner cavity microchannel networks contributes to the low pressure drop. The proposed micro fractal pipette could also innovate the versatile and miniaturized liquid handling system promising to various biological or clinical applications.

Miniaturization of intelligent systems such as lab-ona-chip which integrates several microelectromechanical components allows highly resolved, rapid, safe and lowcost diagnostics with small volume of specimens 1 . However the diagnostics of clinical specimens require additional steps of liquid handling such as collection, preparations, and injection of specimens to biosensors. In biological laboratories, robotic pipetting systems are widely used by automating tedious liquid handling steps but they are limited to be used inside laboratory. To extend the applications out of laboratory, a portable microfluidic sample handling system is inevitable. Conventional syringe and needle require to be tethered by external syringe pump and tubing for precise liquid volume handling. This makes the entire system becomes bulky with several disadvantages such as high priming/dead/residual volumes which are limiting factors towards remote and microfluidic sample handling. Another alternative technology could be the porous micro needles 2 . Although this shows some advanced features such as low dead volume and high portability, the liquid handling performance strongly depends on the geometrical parameters such as the wettable surface area, the porosity and the in/outlet aperture size. Moreover it is essential to control the cavity geometries for the repeatable liquid handling. Most of the porous micro needles have limited porosity and their cavity dimensions are less controlled 2 . Our aim is to propose a microfluidic specimen handling system with enhanced porosity towards biomedical applications requiring specimen/drug collection and delivery. We propose the liquid handling system consisting of highly porous and bioinspired design a) Electronic mail: gilgueng.hwang@c2n.upsaclay.fr; Electronic mail: hwang@if.t.u-tokyo.ac.jp with fractal pyramid microstructure. The fabrication is realized by two-photon absorption 3D nanolithography having sub-micron fabrication resolution. The proposed micro fractal pipette (MFP) allows enhanced porosity, light weight and less invasive liquid handling. Moreover the fractal structures allow facile development of photoresist for inner cavity channels after photopolymerization because the developer medium has better access through large in/outlet openings and also biomimetic microchannel architecture. The MFP does not require external syringe pump and the dead volume of the proposed MFP is close to zero in theory. The liquid filling in and releasing out of the MFP are made by capillary force of the MFP and the target to deliver. This capillary force driven liquid handling is facilitated by low pressure drop (less than 20% from input pressure) thanks to the biomimetically distributed inner cavity channel network. This elim-inates the needs for external pump and additional tube connections for it which are the major source of dead volume for fluidic system. Considering that the MFP is a stand-alone liquid handling system containing exclusively the inner fluidic storage, it can realize near zero dead volume. Similar fractal microstructures were fabricated for surface wettability study 3,4 or for gas permeation membrane 5 . However to the best of our knowledge, our approach of applying the fractal microstructure for the micro pipette is unique. Here we present the MFP's design, fabrication and the liquid handling performances.

Figure 1 shows the schematic concept of the proposed system. First, the MFP is fixed at the end of the capillary probe and the entire system is mounted on the microstage. It should be noted here that the used capillary tube was only fixed to the closed bottom part of the MFP thus the absorbed liquid remains exclusively inside the MFP. Then the petri-dish containing the 2% hydrogel substrate and the red colored ink droplet was mounted onto another microstage. The MFP approaches to the ink droplet to fill itself by capillary force. Our assumption of using MFP is to increase absorption efficiency thanks to the increased in/outlet openings, increased wettable surface area through the cavity and the increased cavity volume. Once the ink is filled, the MFP is moved onto the target hydrogel substrate by stage motion. The inkfilled MFP is approached to get in contact with hydrogel substrate. The filled ink is released onto the hydrogel substrate due to the higher capillary force due to the nanometric scale porosity of hydrogel. Several more attempts can be made onto other position of the hydrogel substrate to completely release the ink out of the MFP. The capillary filling and releasing of the MFP can be understood by the simplified Lucas-Washburn model 6 :

x 2 = γRcosθ 2η
• t by ignoring inertial forces, where x is the liquid filled up distance, η is the viscosity of the liquid, R is the mean capillary radius, t is the time and θ is the contact angle. The contact angle θ is assumed to be static, the capillary liquid motion is proportional to the square root of the mean capillary radius R. Considering the estimated R M F P as around 10.7µm and the sub-microns scale pore sizes of hydrogel, the capillary releasing is made to the hydrogel substrate. Further analyses were made by finite element method simulation to understand the pressure drop of the MFP during liquid handling with the assumption of the laminar flow condition and a static solution. The pressure drop of the MFP was remained to be around 20% which is reasonable value at an input pressure of 100kPa. This can be explained by the fact that the MFP has the inner cavities fluidic channels are close to the biological vascular networks 7 .

To prove the liquid handling by experiments, we designed and fabricated the MFP prototype by two-photon 3D nanolithography system (Nanoscribe, GmbH). This multiscale fabrication capability is essential to fabricate the sub-millimetre scale MFP prototype. After photopolymerization process, the fractal structure with large openings with the connected cavities allowed facile access of the developer medium to the inner cavities for the development of photoresist. The developed structures were rinsed several times by isopropyl alcohol and dried carefully. Scanning electron microscope imaging was made to observe the quality of the process. As shown in the figure 1, the MFP was fabricated with good quality of inner cavity development. For the target substrate, 2% agarose hydrogel samples was prepared. For the scaledup MFP, polylactic acid (PLA) was 3D printed to have young's modulus around 3.5GPa similar to the epoxy resist (around 4GPa). The MFP was scaled up by 55 times thus the width of the base part is 55mm.

The fabricated MFP was applied to demonstrate the liquid handling performance. The performance of liquid handling can be understood by the efficient absorption from source, handling and release to target with minimal loss. Since the liquid handling typically comprises several inefficiencies at each steps, liquid handling performance tests were made by following the schematic concept shown in the figure 1. As shown in the figure 2, the MFP mounted at the end of capillary probe was operated by micro stages. The MFP entered the red dye droplet to fill it up. Once it is filled, it moves out of the droplet. The MFP's gradually decreasing inner cavities allow the filled liquid to remain inside the cavities. The surface area of whole cavities inside single MFP is 8.5 times larger than full pyramid. The filled liquid through connected cavity channels could avoid rapid evaporation during handling. Then the filled liquid was released to target hydrogel substrate by stage motion. The MFP demonstrated to be able to release several spots by repeating the process without getting dried. After this process, we can confirm that the red dye filled MFP was completely drained out and released to the hydrogel. We estimate the filled liquid volume is around 185 nl which can obtain the porosity around 78%. This contributed to the 28 times higher absorbability of the MFP compared to the full pyramid. Furthermore the penetration of the MFP through the hydrogel was guaranteed thanks to the solid mechanics from the photopolymerized epoxy based resist having young's modulus around 4GPa superior to the target hydrogel substrate (2% agarose) having the young's modulus in the range of 14 kPa 8 . If the struc-ture has lower young's modulus compared to the hydrogel target substrate, the absorbed liquid in the pipette could be lost at the interface of the hydrogel substrate due to the deflection by compression during the insertion. Therefore as in the proposed case, forming the MFP with hard epoxy avoids the loss of the liquid during insertion process. This can also be compatible to the human's skins having large range of young's modulus 9 . The very performing liquid handling of the MFP is also partly supported by the large surface opening area (1.32mm 2 thus 76% of 4 sides of pyramid) of the MFP allowed facile liquid absorption and release. This can be explained by the fact that there are large openings at the contact surface of MFP which reduces the fluidic resistance at the interface to the target hydrogel substrate. In comparison with the conventional micro porous structures, there are much less chance to get also clogged in the openings. Mechanical resistance of insertion to hydrogel is also important for the successful liquid delivery. Highly resistive pipettes could lead to the application of excessive insertion force thus the insertion could increase the pain or the incision. The proposed MFP has only 0.42mm 2 (24%) of the outer surface in contact with hydrogel thus the friction is reduced compared to the full pyramidal shape. We suppose that this reduced contact surface thus the insertion friction can contribute to reduce the insertion pain or incision. The reduced contact surface area also reduces the liquid loss at the outer surface of the pipette during insertion. Together with small gravity force, the MFP can be simply fixed to capillary tube by double side adhesive film. Large capacity of liquid absorption by single MFP reaching up to 185nl which is 78% of the volume total of pyramid could cause eventual loss of liquid during displacement step as well due to the gravity. During the displacement step, the liquid filled MFP contained the liquid without any loss. This is thanks to the 8.5 times larger surface area in contact with the filled liquid. This increases drastically the surface capillary force which contributes to hold the liquid stably during displacement even with mechanical vibrations. The minimum spot size achievable with the current dimension of the MFP could be as small as below 10 µm in radius which can be estimated from the minimum deliverable volume of liquid is 3.8 pl. Considering the smallest the cavity dimension of the tetrahedron is with the side of 20 µm and the sub-micron resolution 3D nanolithography, further miniaturization could be attempted. To test the scalability of the concept, we used the 3D printed scaledup MFP. After making the same steps shown in the figure 1, the absorbed volume of liquid (10 ml) was transported without any loss during handling. Considering the fact that we did not make Reynolds number matching notably the viscosity of the liquid in the scaled-up model tests, stable liquid handling is highly confirmed.

The proposed MFPs demonstrated several enhanced features compared to the conventional approaches. First, they can fill up large volume percentage of liquid up to 185nl which is 78% of full pyramid thanks to enhanced porosity. Second, thanks to the increased contact surface area and stable mechanics, the filled-up liquid was able to be delivered stably to target hydrogel substrate without showing any loss during the displacement and also insertion steps. As summarized in the figure 3, the proposed concept also demonstrated to be effective in the liquid handling scalable volume from picolitre to millilitre without any support from external microfluidic pumps or syringe. Compared to porous microneedle with equivalent porosity, the MFP shows advantages in the lower mechanical resistance of insertion and fluidic resistance thanks to biomimetic inner cavity channel network promising to drug delivery through the skin or other tissues. The MFP is easily washable for more than hundreds of times of repeatable use and did not show ageing effect at least for few months so it is theoretically semipermanent in its lifetime. Considering also the miniaturization and the ignorable dead volume feature, the MFP could be promising to portable liquid handling tool thus the applications could range from the field specimen collection to the liquid handling inside laboratory.

The authors would like to acknowledge Gwenaelle Bonfante and Hakjae Lee for agarose prepartion. The work was partly supported by the french RENATECH network of CNRS for 3D nanolithography, the MEXT nanotechnology platform of the University of Tokyo and also by the JSPS Core-to-Core Program (A. Advanced Research Networks).

The data that support this study are available from the corresponding author upon reasonable request.

FIG. 1 .

 1 FIG. 1. (Top) Proposed concept of the MFP for liquid handling. The MFP approaches to fill the liquid by capillary force and it moves to the target (2% agarose gel) to release the liquid by capillary force of hydrogel. (bottom) Design and the fabrication of MFP. The fractal pyramid 5 levels with 1mm 2 base was fabricated by two-photon 3D lithography.

FIG. 2 .

 2 FIG. 2. (Left) Liquid handling experiments by MFP to hydrogel. (Right) Performance comparison between full pyramid and the MFP. The scale bar is 3mm.

FIG. 3 .

 3 FIG. 3. Comparisons with classical needle, porous microneedle and the proposed MFP.