The microjet injector system accelerates drugs and delivers them without a needle, which is shown to overcome the weaknesses of existing jet injectors. A significant increase in the delivered dose of drugs is reported with multiple pulses of laser beam at lower laser energy than was previously used in a Nd:YAG system. The new injection scheme uses the beam wavelength best absorbable by water at a longer pulse mode for elongated microjet penetration into a skin target. A 2.9 μm Er:YAG laser at 250 μs pulse duration is used for fluorescent staining of guinea pig skin and for injection controllability study. Hydrodynamic theory confirms the nozzle exit jet velocity obtained by the present microjet system.
© 2012 Optical Society of America
The laser-based microjet injection system uses the hydrodynamic impact of a narrow liquid jet onto skin. The immediate delivery would enable minimized prescription of topical drugs intended to work on the outer layers of the skin, avoiding any skin irritation or allergic reaction and preventing uncontrolled evaporation of active ingredient and unpleasant odor associated with noninvasive procedures. Several types of injection mechanism have been considered, including spring compression, expansion of piezoelectric transducer, linear Lorentz-force-driven piston actuator, and expansion of laser-initiated waves in water [1–3]. Such mechanisms would eliminate abundant needle wastes, and they are favored for highly needle-phobic patients .
A narrow, high-pressure jet of 100 to velocity is required to accelerate the drug to penetrate the animal skin with of yield strength . Reducing the jet diameter to a 100 μm size has shown the advantage in drug delivery of minimizing damage to the tissue . The present scheme of injection via Er:YAG laser beam at 250 μs pulse duration generates pressure by the displacement of liquid via laser-induced vapor bubbles and the elastic pumping of the drug through a nozzle by a membrane separating the driving liquid from the drug.
Success in jet injection requires sufficient impulse of the jet to penetrate the target tissue. In the case of an Nd:YAG laser at 7 ns pulse duration, high irradiance of a -switched beam produced instantaneous expansion of bubbles and generated multiple shockwaves . The consequent jets due to both shock and bubble expansion reached while the short length of the jet is insufficient for ensuring a sizable dose of drug targeting a treatment site.
Here we adopted a quasi-long-pulsed Er:YAG beam at 250 μs to source an injector. The high absorbance of the beam to water (2940 nm) also enables a stronger generation of a jet at lower laser energies. We present the enhanced controllability and dosage of delivered drug into a guinea pig’s skin through fluorescent staining on both postmortem abdominal and living dorsal skins. Furthermore, to verify test results, the Plesset equation of the vapor bubble theory has been adapted to confirm the measured jet velocity resulting from the laser-initiated microjet.
When the low-irradiance laser energy reaches the driving liquid in the upper chamber of the injector , the temperature rise in a focal point in water admits a sudden vaporization and generation of a vapor bubble . The process of constant-pressure phase transformation is marked by a superheated liquid formation of vapor voids. The Er:YAG laser has 250 μs pulse duration and wavelength of 2940 nm at which water best absorbs the beamed energy. Once vapor is accumulated in the bubble, any additional radiation passes through the expanding bubble without being directly absorbed by the vapor. Elongation of the bubble discontinues as the initiation of the bubble stabilizes beyond 250 μs and reaches a maximum radius.
We adopt explosive bubble expansion induced by a laser irradiation as an actuator for ejecting a coherent microjet. The present injector consists of a micronozzle for storing liquid drug, a chamber for driving fluid separated by a heat-resistant flexible membrane between drug and water, and air-tight confinement glass at the beam incident end with O-ring-type sealing. A highly water-absorbant Er:YAG beam irradiates the water for vapor bubble generation within the driving liquid chamber. Ideally sealed in the chamber, growth of the bubble would cause a sizable pressure impulse on the elastic membrane. The elastic response of the membrane ejects liquid drug out from a 150 μm nozzle at a velocity needed to penetrate skin.
In Fig. 1(a), the ejected jet shown in an air background reaches , which has larger injection volume than the speed previously attainable by a nanosecond pulsed Nd:YAG laser system . The laser system offers various jet properties with a change in laser energy. The flow through a narrow nozzle experiences turbulent and frictional losses. Figure 1(b) indicates that friction at the nozzle exit causes velocity reduction in the region below 800 mJ and the turbulence shortens the jet length in the region beyond 800 mJ . The instability causes spray and a decrease in energy of the jets, a process known as atomization. In all injections of the present scheme, the standoff distance between the nozzle and the skin is less than 3 mm to avoid any instability due to these jet properties.
The injection liquid was prepared by dissolving fluorescein isothiocyanate (FITC, ) in dimethyl sulfoxide (DMSO) solution for verifying microjet penetration performance. The treated skin from the FITC test was analyzed by a fluorescence microscope (Nikon Eclipse Ti-U), and hematoxylin and eosin (H&E) staining was used to monitor alteration of tissue after injection.
The stained skin samples were frozen and chopped by cryotome (Leica). Three pieces of a sample were made with the injection point being the center. Each cross section was vertical to the injected area and made into a slide to analyze with a microscope. The embedded sites of fluorescent trace observed with the microscope would confirm drug penetration and the range of spread underneath the outer skin layer.
A male Hartley guinea pig was laid on a table to perform jet injection on the back for a dorsal skin test. In the abdominal case, the relevant skin was first sectioned into a target and cleaned to remove fat and subdermal tissues. The prepared skin sample was affixed on a foam board with pins. The laser irradiated the driving fluid, and the jet was ejected from the nozzle vertically to a target material. The guinea pig used was 250 grams of weight. The skin was epilated with wax for 24 h before the experiment and immediately used without freezing.
For the skin mockup for instant visualization of penetration, a gelatin gel was used, which offers controllable mechanical properties depending on its weight percent. Gelatin of 60 Bloom was dissolved in water at 5 percent. The Bloom number indicates the toughness of gels, and 60 meets the typical animal skin toughness.
Figure 2 is the staining result on a sectioned sample of guinea pig abdominal skin. The fluorescence is dispersed in all directions around an impact point. The microjet initiated with a 19 J laser beam delivers the drug over the epidermis and dermis within from the skin (Fig. 3).
Figure 3 shows the results of microjet injection on dorsal skin as administered on a living guinea pig laid on a table. The drug is evenly dispersed on the skin tissue similar to the abdominal case. The dorsal skin is thicker () than the abdominal skin such that deeper wetting of the relevant layers is effectively treated with FITC. The H&E staining shown in the upper-right windows after microjet injection displays the architecture of the treated sites, and it shows that the drug is delivered with no alteration of skin morphology adjacent to the injection site. The microjet initiated with a 1.57 J laser beam, however, achieves a targeted local delivery rather than dispersion [Fig. 3(b)]. Unlike the abdominal case of a sectioned sample affixed to an acrylic plate, the underlying structure of dorsal skin supports jet propagation, allowing a deeper penetration. Even though the injection spot had been ruined following the path of a jet, these microstructural changes are expected to be recovered by the barrier recovery process .
The jet produced with 1.19 J of laser pulse showed smaller volume than the case of 1.57 J. The jet energy is mostly dissipated at the upper layer of skin. With higher energy, however, increased jet energy reaches deeper layers of skin. The jet energy is converted to deformation of the skin barrier or propagation of the stress wave depending on the skin properties. The jet achieves farther delivery of drugs when more energy is converted to deformation energy than to stress wave.
To evaluate jet injection, gel models are used for visualization and mimicking of the skin elasticity. Jet injection consists of three events: jet impingement, flow into skin, and dispersion under skin [Fig. 4(a)]. Jet impact creates a hole on the gel with an estimated impact pressure or water-hammer pressure from Eq. (1), where the impact pressure depends on the sound velocity as well as the jet velocity . Then at a lower jet pressure proportional to the square of the velocity, the ejected dose is delivered into the gel making a path of jet stream. A thin cylindrical jet of 150 μm generated from an injector causes virtually zero splashback at the contact surface as seen in Fig. 4(a), allowing smooth penetration. In gel models, the dense structure with no porosity forces the jet to agglomerate and causes bounce of the drug in the gel:1) suggests a minimum jet velocity of for water density and sound speed of in water.
The penetrated depth and injected volume are evaluated with varying laser energy on a gel model [Fig. 4(b)]. The penetrated depth increases as the laser energy is intensified up to 800 mJ. At 600 mJ, the injection efficiency is reduced due to the recovery response of the gel. Above 800 mJ, increased jet velocity shortens the jet length according to the jet breakup. The spray characteristic and instability of the weakening jet beyond a high critical Weber number may be responsible for this observation.
The jet velocity may be analytically determined. We consider a 408 mJ (low-energy) case for illustration purposes. We used the empirical data for temporal development of the bubble radius and the Rayleigh–Plesset approximation  to estimate the pressure and temperature gradients on the driving liquid wall for jet ejection. For vapor bubbles, the thermal effects play a dominant role and the effect of liquid inertia can be neglected. The Plesset model considers the evaporation and the heat conduction. The temporal evolution of bubble radius can be approximated by the following equation 
Solid circles in Fig. 5(a) illustrate the empirical Our empirical data obtained from the images of 5(b) implies that . The dependence of Eq. (2) for this value of is illustrated with the solid curve. A good agreement between asymptotical behavior and the empirical data is observed. Equations (1) and (2) suggest that the “superheat” should be nearly constant, and by substituting values of the dimensionless radius, , , and for water, we estimate .4), we estimate and thus .
Suppose that the chamber is closed and the volume of the liquid is comparable with the actual volume of the bubble in the driving chamber of the injector. Then in Eq. (4) should decrease, taking into account the compression of the liquid inside of the chamber. This is done by the Tait equation of state,
From Eq. (5), the pressure on the chamber wall increases with the radius of the bubble. The bubble grows until the pressure in the liquid compensates the pressure on the bubble wall. Assuming that the bubble has the same initial value of , the maximum pressure on the chamber wall should be reasonably close to . The rise of the pressure inside of the chamber pushes the liquid out of the chamber. The resulting jet velocity can be estimated with the help of the Bernoulli equation,2(b).
In this Letter, the performance of an Er:YAG laser-initiated microjet is evaluated for transdermal drug delivery. Approximately 500 nl per pulse of drug is delivered beyond the skin barrier in the form of a microjet as the injected drugs are effectively dispersed over the epidermis. We ensured controllability of the laser-initiated microjets via the longer pulse (250 μs) at lower energy (), which is an improvement from the Nd:YAG-based injection scheme delivering 200 nl per pulse .
We thank the Korea Research Foundation (DOYAK-2010) for financial support through the Institute of Advanced Aerospace Technology (IAAT) at Seoul National University.
1. A. Arora, M. R. Prausnitz, and S. Mitragotri, Int. J. Pharm. 364, 227 (2008). [CrossRef]
2. T. Han and J. J. Yoh, J. Appl. Phys. 107, 103 (2010). [CrossRef]
3. A. Taberner, N. C. Hogan, and I. W. Hunter, “Needle-free jet injection using real-time controlled linear Lorenz-force actuators,” Med. Eng. Phys. (to be published). [CrossRef]
4. S. Mitragotri, Nat. Rev. Drug Discov. 5, 543 (2006). [CrossRef]
5. C. E. Brennen, Cavitation and Bubble Dynamics (Oxford University, 1995).
6. A. H. Lefebvre, Atomization and Sprays (Hemisphere, 1989).
7. J. A. Segre, J. Clin. Invest. 116, 1150 (2006). [CrossRef]
8. M. S. Plesset and A. Prosperetti, Annu. Rev. Fluid Mech. 9, 145 (1977). [CrossRef]