Needles are a common and efficient method for drug delivery, used for more than two centuries. However, it presents serious health issues, such as waste contamination, risk of spreading diseases, unwanted needle-stick incidences, pain, and phobia, among others.1 The reuse of needles is a common practice, which may transmit contagious diseases and deadly viruses. In Africa, it is estimated that 20 million injections contaminated with blood from HIV-infected patients are administered inadvertently every year.2 In addition, according to the World Health Organization, unsafe injections in the year 2000 will lead to 9 million deaths between 2000 and 2030.3
Any candidate device to replace the widely used needles must be cheap, portable, easy to operate, and safe. Such a device should be reusable, no physical contact should take place between the device and the skin of the patient, reducing contamination risks. Different mechanisms for needle-free jet injection have been investigated.4–21 These can be classified as impulsive pressure-induced jets, such as compressed gas, spring6,7 or piezoelectric transducer,8–10 and cavitation-induced jets, such as electric current11 or laser.17–21 Particularly, spring and compressed gas systems are now commercially available, which are mostly used for insulin injection.13–16
The potential of cavitation-induced jets by lasers was first explored by studying cavitation bubbles near to an elastic boundary.22,23 Laser-based systems produce extremely fast jets (up to ) and reach an injection depth of 5 mm,24 which is sufficient for most medicines, and can operate without cross-contamination. In this case, pulsed lasers are used to generate plasma or vapor bubbles which, in turn, emit pressure waves of several GPa of amplitude.25,26 The required pulsed lasers are expensive, noisy, and heavy. Recently, we demonstrated a jet injector, which is based on continuous-wave (CW) laser cavitation27 (or thermocavitation28–31). These lasers are widely used and have the benefits of being lightweight, cheap, and pose a much lower safety risk than pulsed lasers for a given amount of light energy, as the irradiance is lower. However, the injection velocities were still limited to , which is barely sufficient to penetrate the outer skin layer (stratum corneum, which has a thickness of 10 to ),32 and the injection depth was not assessed.
In this work, we present a microfluidic device that resolves this issue. A tapered shaped channel enhances the jet velocity in combination with dynamic focusing. Experiments of liquid jet penetration into agarose gel at 1% were also performed as a proof of concept for an eventual use of these designs in injection devices. CW lasers may solve the problem of integrability in portable devices since they are compact, cheap, and powerful enough to achieve jet speeds able to penetrate skin. Finally, we present an overview of different jet injection methods and discuss current applications of laser-based jet injection as well as future avenues toward the improvement of CW laser-based systems.
Materials and Methods
The setup depicted in Fig. 1(a) was used to study the bubble dynamics and the corresponding liquid jet propagation. Thermocavitation was produced by light absorption in a copper nitrate saturated solution. The laser () was focused at the bottom of the device chamber with a microscope objective. The spot produced an intensity of ( and beam waist of ). The microfluidic device was placed on a linear translation stage holder in order to align it with respect to the laser spot. A fast camera (FASTCAM SA-X2) and a white-light source (Olympus LP-1) were placed at different locations to study two different events: (1) jet propagation (velocity and shape) and (2) liquid penetration into agarose gel, as indicated in Fig. 1(a). These events were recorded at 300,000 frames-per-second (fps) and 125,000 fps, respectively. The laser current was controlled with square wave signal from a function generator, which also triggered the fast camera. The laser was turned on for 500 ms, which allows to observe a full cycle of bubble expansion and collapse. If the laser current is not modulated, then a quasiperiodic bubble formation, whose frequency is controlled with the laser intensity, takes place.30
Description of the Device
Microfluidic chips were designed and fabricated in glass substrates under clean-room conditions. Two wafers of Boroflat glass were identically micromachined with wet-etching in hydrogen fluoride solutions, and then placed together with anodic bonding. The bonded wafers were then diced in chips with dimensions of . The chips were constituted by (i) a channel inlet ( depth); (ii) a circular container ( depth); (iii) an s-shaped channel ( depth), to provide fluid resistance and control over the liquid volume inside the device; (iv) a chamber ( depth), where the cavitation bubble is created; and (v) a straight or tapered channel ( depth) for liquid propagation and confinement, as is shown in Fig. 1(b).
The device geometrical parameters that affect the liquid jet velocity are the channel diameter at the exit , channel diameter before the tapering , taper ratio , and chamber length , as shown in Fig. 1(b). Several devices were fabricated with dimensions of , 200, 300, and , , where is two times the etched depth, and ranges from 200 to . Although the cross-section of the channel was not axisymmetric, it has been demonstrated that for asymmetric nozzles, the jets spread only slightly faster at subsonic conditions (),33 which is our case. Hence, the generated jets in this study can be assumed as cylindrical.
In previous reports, it was found that the cavitation bubble expansion rate is proportional to the chamber width.27,34 For this reason, this parameter was set to , which is sufficiently large to allow a fast expansion, and on the other hand, not too large as to waste its kinetic energy in displacing a large liquid volume, which may slow down the jet speed.
The channel length was set at , however, it extended beyond the outlet of the device, as shown in Fig. 1(b). This was done to prevent changes in the taper ratio values previously established, due to the lack of precision in the cutting and separation process of individual devices from the wafer. The liquid inlet was connected to a 1-mL plastic syringe Terumo (Terumo medical products) through a glass capillary tube with diameter, using a microfluidic fitting and a connector. The syringe was used to manually control the position of the meniscus in the channel by changing the liquid volume inside the cavity. A saturated solution of copper nitrate (13.78 g in 10 mL of water), with an optical absorption coefficient of for the laser wavelength, was used to produce the cavitation bubble.29
Agarose gel (Sigma-Aldrich) at 1% was prepared in small cubes of approximately, to characterize the penetration depth of the liquid jets produced by the devices. It has been proved that agarose gel is an appropriate model to compare with human skin,35 since its mechanical properties are similar to soft tissue in the body.36 However, it is important to mention that human skin, specially, the stratum corneum (outermost layer of skin) is a very complex tissue and its properties change significantly depending on the part of the body. The Young’s modulus of skin can have values ranging from 20 kPa to 2 MPa, depending on the part of the body and per individual (age, hydration level, and many other characteristics).32,37 The Young modulus and plastic yield stress of agarose at 1% are around 4038 and 30 kPa,39 respectively, which lies on the lower limit of the skin Young’s modulus.
Position of the Meniscus
The presence of a concave liquid–air interface plays a crucial role on achieving high speed liquid jets.17 In fact, the dynamic focusing, driven by the pressure wave produced by optical breaking, is due to the initial contact angle of the liquid with the channel walls. The meniscus concavity can be tuned using surfactants to reduce or increase the contact angle.17,40,41 For large contact angles (), the jet speed is minimum but increases as the contact angle is reduced. We confirm these findings, but, since the device configuration is different from capillary tubes used before,17 we have investigated this effect further.
The parameters that characterize the liquid–air interface are shown in Fig. 1(c), the distance between the laser focus (where the bubble is created) and the meniscus position is , whereas is the initial contact angle of the meniscus. The initial contact angle was obtained with the relation:17
The initial radius of curvature is varied by minor adjustments of the liquid volume with the syringe. The three cases under study were (i.e., the meniscus formed is rather small and the contact angle is ), meniscus (the device filled until half of the channel), and (only the chamber device is filled), as shown in Figs. 1(c)–1(e), respectively. The values of and for each case in the device with parameters , , and are presented in Table 1.
Values of the parameters describing the meniscus characteristics and the corresponding jet velocity for device with Dx=120 μm, n=0.5 and E=200 μm.
Jet Velocity Parametric Study
Figure 2 shows a typical example of the jets produced by the fast expanding bubble in a device with the following parameters: , , and laser intensity , for the meniscus , , and . The bubble drives the meniscus dynamic through the channel, leading to dynamic focusing of the flow and finally to a jet. For , the jet has a tip around five times smaller than the rest of its body, with a speed up to during the first , while and , only and , respectively.
With these image sequences, we can observe that for (), a sharp jet is formed. When increases, the liquid jet becomes less focused and its sharpness decreases (). Finally, when no meniscus is present (, ), there is no focusing and therefore, the jet becomes blunt, due to the liquid adhesion to the walls of the output channel. Under these circumstances, the fluid is pushed outside of the device with a tip bigger than the rest of the jet.
The jet velocity , as a function of the taper ratio , for a channel diameter , chamber length , and laser intensity of , is shown in Fig. 3(a). As expected, the jet velocity increases as the taper ratio is reduced. However, the meniscus position and shape have a substantial effect on the jet velocity. For and , increase on the velocity was achieved when decreased from 1 to 0.25. For , the jet velocity is practically independent of the taper ratio . This result highlights the importance of dynamic focusing to achieve high speed jets. It is expected that flow through a tapered channel increases the velocity according to the principle of continuity as , where is the velocity inside the channel with diameter .42 Thus, for taper ratio of 0.25, the jet velocity increases by a factor of 16. This finding is in disagreement with the mere 1.6 ratio observed in the experiments [see Fig. 3(a)]. The following reasons are attributed to this mismatch. The chamber is not filled completely, and the liquid is displaced by a half hemisphere shaped bubble and not by a constant plane as in the Bernoulli equation. Then, for the latest case, the exerted pressure is constant, whereas for our case, there is a pressure gradient. Since the channel is partially empty, friction may also play an important role reducing the jet velocity.
The dependence of the chamber length on the jet velocity is shown in Fig. 3(b) for , , and . Note that the fastest jets were obtained for the smallest length () regardless of the initial contact angle value. These results are in good agreement with previous studies in capillary tubes, where it was found that the jet velocity is inversely proportional to .17 In addition, the focused jet is no longer in contact with the channel walls, hence, kinematic friction is reduced. These results confirm the relevance of dynamic focusing for . In capillary tubes, where the walls are parallel, a smaller contact angle results in faster jets. However, in tapered channels, it is not necessarily true. Larger contact angles, in combination with the device geometry, can lead to more efficient focusing, as is shown here.
The jet velocity as a function of channel diameter for different taper ratios is shown in Fig. 3(c). The jet velocity increases as both the channel diameter and the taper ratio are reduced. However, for diameters , the velocity remains almost independent of the taper ratio. As the diameter increases, the fluid confinement is reduced and the liquid is less focused, giving as a result the reduction of the jet speed to a minimum value. Based on the results shown, a maximum jet velocity up to was observed for the following parameters: , , , and ( and ).
The initial conditions of the meniscus not only influence the jet velocity but its shape, too. A meniscus as near as possible to the bubble formation place and a small contact angle are the conditions that will lead to a fast and sharp jet. As was observed before, in order to increase the jet velocity, the taper ratio, chamber length, and channel diameter need to be reduced. However, the reduction of these parameters will eventually lead to an early breakup of the jet into small droplets, changing from a jet regime to spray regime.43
Skin Phantom Penetration
The performance of the fabricated devices to generate jets was tested for skin phantom penetration. The penetration depth into agarose 1% gel cubes as a function of channel diameter and jet velocity for meniscus was measured. The velocities obtained for ( to ) are not high enough to obtain a large penetration into agarose, whereas in the case of ( to ), the thickness of the jet tip is too small, so it was difficult to observe the real penetration depth in the gel with the fast camera. Thus, meniscus was chosen since the jet velocities ( to ) are sufficiently high.
A jet expelled from the channel, and breaking through the gel with a velocity of , is shown in the image sequence of Fig. 4(a). In this particular case, the jet penetrated a maximum distance of in . The penetration depth , as a function of time, is shown in Fig. 4(b). These values were calculated from image sequences, as presented in Fig. 4(a). A data fitting curve shows a behavior of the form . The exponential behavior was predicted in Ref. 24.
The penetration depth as a function of the jet velocity for and is shown in Fig. 4(c). As expected, the penetration depth increases as the jet velocity increases. A maximum penetration depth up to was achieved for and . The parameters studied in the microfluidic devices and how they affect the penetration depth of the liquid jet into agarose are shown in Table 2. It can be observed that by reducing the parameters of the device , , and , and by reducing the liquid–air interface parameters and , the penetration depth will reach its maximum.
Penetration depths L dependence on the geometrical parameters of the microdevice, and parameters of the liquid–air interface. The up arrow and down arrow indicate an increment and a reduction on those parameters, respectively. The number of circles in L, represent the penetration length.
|Channel diameter Dx (120 to 500 μm)||Taper ratio n (0.25 to 1)||Chamber length E (200 to 1000 μm)||Contact angle θc (47 deg to 90 deg)||Bubble distance B (200 to 700 μm)||Penetration L (0.390 to 1 mm)|
The liquid volume injected in the gel was calculated from the initial liquid volume contained inside the device, and subtracting the remaining liquid once the jet is expelled, assuming evaporation is minimal. For , were introduced into the agarose, whereas for , a volume of 157 nL was delivered.
An overview of jet velocity measurements for both pulsed-laser and our CW laser system is provided in Table 3. In both laser configurations (pulsed and CW), a pressure wave is created by vaporizing a small amount of liquid. However, in thermocavitation (CW laser), the intensity threshold to produce bubble nucleation is several orders of magnitude smaller than for pulsed systems, as a sufficient amount of energy can be delivered over a longer time span. For example, a pulsed laser intensity around was necessary to generate jets of in previous work,17 whereas for our CW system, an energy of was required.
Laser characteristics of recent investigations in jet injection systems and maximum jet velocity achieved.
|Laser||τp (ns)||E(mJ)/I(W/cm2)||Vmax (m/s)||Ref.|
|Nd: YAG 1064 nm||5.5||1400/—||200||44|
|Nd:YAG 532 nm||5–9||100/—||264||20|
|Nd: YAG 532 nm||6||850||17|
|Er:YAG 2940 nm||2500||1000/—||45||18|
|Nd:YAG 532 nm||6||250||24|
|Nd: YAG 1064 nm- Er: YAG 2940 nm||7-2500||408/—||30–80||19|
|Infrared laser 790 nm||—||—/||30||27|
|Infrared laser 790 nm||—||—/||94||This work|
Remarkably, for CW lasers, the jet velocity decreases with increasing laser intensities, in contrast to results for pulsed lasers.17 This is due to the fact that the spinodal limit is achieved faster as the laser intensity is increased. Therefore, for high laser intensities, thermal confinement is achieved faster, limiting the bubble size, whereas for lower intensity, heat diffusion allows greater superheated volume and therefore bigger vapor bubbles.27,29 Jets generated with CW lasers may therefore have a lower maximal velocity compared to pulsed lasers, but the current work shows that these velocities are still sufficient for injection purposes.
Common injected drugs include antibiotics, steroids, hormones, vaccines, and insulin, among others. Typical dosage volume and injection depth of these medicines are plotted in Fig. 5(a). In addition, the volume injected, per injection event, as a function of penetration depth, for different jet injection methods, is shown in Fig. 5(b). It can be observed that for impulsive pressure-induced jet systems (spring, gas, chemical reaction, and piezoelectric actuator), the injected volume reaches medical doses like insulin, vaccines, and antibiotics. By contrast, for cavitation-induced systems (electric current, pulsed laser, and CW laser), the maximum volume achieved () is still below the medical doses.
The liquid quantities injected by these laser systems are too small for typical drug doses still. Repeated injection may offer a solution to this problem, as cavitation repetition rate of 4 kHz could be reached,29,30 so a typical dose of 1 ml could be achieved in a few seconds. Further investigations from a medical perspective would be required to validate this approach.66 As shown in Fig. 5(a), applications where the volumes required are smaller, such as allergy tests,67 medical tattooing,57 and microdosing for clinical study68–70 could be achieved with CW laser-based injection.
Another important issue to further investigate in our device is the injection of real drugs for medical treatment, instead of the copper nitrate solution we have used in this investigation. One option could be to use different laser wavelengths according to the absorption coefficient of the drug, however, this will probably affect the chemistry of the injected solution in an undesired way. In order to avoid this, it is necessary to thermally isolate the cavitating liquid from the drug one, as proposed elsewhere.20 Additionally, the jetting of liquid drugs with different viscosities needs to be studied for possible implications in jet velocity and skin penetration depth, as was recently investigated with the use of impulsive pressure acceleration for the generation of highly viscous jets.71
The penetration depth values presented in this work are comparable with previously reported results obtained for pulsed laser,24 where a depth of 1 mm was reached for a jet speed of in a capillary tube with of diameter. Earlier studies22,23 have demonstrated jet velocities up to penetrating an elastic boundary, even through a water layer of . Yet another investigation, where cavitation-induced jets were generated by an electric discharge, penetration depths up to were achieved, with jet velocities between 130 and .11 Notwithstanding the higher velocities than those we obtained here, less penetration depth was achieved. This may be attributed to the large scatter nature of the jets.
According to the literature,72 the stratum corneum has a thickness between 10 and , whereas the epidermis is between 40 and , depending on the part of the body. Therefore, by combining CW laser and microfluidics systems, it should be possible to penetrate human skin, opening, for the first time, the possibility of cost-effective, portable, and silent jet injection technologies for intra- or transdermal drug delivery.
CW laser-based cavitation is a mechanism to produce jets with sharp shapes and velocities sufficiently high to penetrate the skin. Here, different geometries were studied, and we found a maximum jet velocity of for a channel diameter of , chamber length of , and taper ratio of 0.25.
The dynamics of the liquid–air interface inside the microfluidic device determines the velocity and shape of the jet expelled. The meniscus focuses the liquid inside the channel leading to a fast and sharp curved shape jet. As the taper ratio is reduced, the jet velocity increases and the same for chamber depth. However, if the surface is nearest to the initial bubble and with a contact angle , then the speed remains almost constant with the variation of the geometrical parameters. Penetration depths into agarose 1% gel up to 1 mm were reached.
In this work, the potential of CW-laser based microfluidic systems as needle-free drug injector has been demonstrated. The significant advantage of our proposition is that less energy to produce liquid microjets is required than with pulsed lasers, besides its lower cost and better portability. However, further studies of penetration depth into skin need to be carried out to fully validate this technique. In particular, higher volumes will be required to reap the benefits of CW-laser based jet injection.
The authors have no relevant financial interests in this article and no potential conflicts of interest to disclose.
The authors acknowledge financial support from CONACyT-Mexico and are grateful to Prof. Detlef Lohse and Prof. Chao Sun from POF for their support, advice, and access to their laboratory equipment. Also, a special thanks to Remco Sanders and Jeroen Korterik for laboratory assistance. DFR acknowledges the recognition from the Royal Dutch Society of Sciences (KHMW) that granted the Pieter Langerhuizen Lambertuszoon Fonds, 2016.