Autore:
Angela Longo
Relatori:
Ing. Giuseppe Barillaro Prof. Andrea Nannini
Microneedles for minimally-invasive
transdermal medical applications
Anno 2014
UNIVERSITÀ DI PISA
Scuola di Dottorato in Ingegneria “Leonardo da Vinci”
Corso di Dottorato di Ricerca in
INGEGNERIA DELL’INFORMAZIONE
SOMMARIO
Questa tesi tratta il design, la fabbricazione e la caratterizzazione di dispositivi minimamente invasivi per applicazioni transdermiche di prelievo e rilascio di fluidi. In particolare, i dispositivi proposti permettono di prelevare piccole quantità di fluidi dagli strati superficiali di pelle e rilasciare farmaci per via transdermica.
In primo luogo, viene presentato un nuovo dispositivo per l’auto-monitoraggio del glucosio in pazienti diabetici: tale dispositivo, chiamato glucose pen, si basa su una matrice di microaghi in ossido di silicio, in grado di essere inserita nei primi strati della pelle e, da qui, prelevare una piccola goccia liquido interstiziale. Questa goccia viene testata, grazie ad un sensore di glucosio posto sul retro della matrice di microaghi, in modo da misurare la quantità di glucosio ivi contenuta. Quest’operazione è molto meno invasiva rispetto al classico metodo di auto- monitoraggio del glucosio, poiché riduce il dolore, evita il sanguinamento e, di conseguenza, anche il rischio di infezioni. Verranno descritti design, fabbricazione e caratterizzazione meccanica della matrice di microaghi; inoltre, verrà illustrato il design di tutti i componenti della glucose pen (il circuito di pilotaggio e lettura del sensore di glucosio, l’assemblaggio del sensore di glucosio e della matrice di microaghi in un’unica testina usa e getta, il montaggio della testina nella penna ed il meccanismo di attuazione che permette il movimento della matrice di microaghi ed il suo inserimento nella pelle).
Successivamente, viene presentato un processo basato sul micromachining elettrochimico per la fabbricazione di matrici di microaghi in silicio. Con il processo presentato sono stati realizzati microaghi, appuntiti o piatti, da utilizzare per essere inseriti nella pelle e, in questo modo, migliorare il rilascio transdermico di farmaci, che possono permeare attraverso i microcanali creati dai microaghi stessi nella pelle.
Infine, viene presentata l’attività svolta presso il “Dimes Technology Center”, Delft (NL), per la realizzazione di matrici di microaghi in silicio, connessi con un serbatoio posto sul retro della matrice stessa. Il processo sviluppato si basa sul Deep Reactive Ion Etching e viene effettuato sul wafer di silicio intero (e non sul singolo dispositivo), permettendo così di realizzare matrici di aghi molto diverse in termini di forma, dimensioni trasversali e densità con un unico processo. Il processo sviluppato si configura come una valida alternativa per realizzare a livello industriale microaghi per rilascio di farmaci e prelievo di fluidi, poiché è wafer level e sfrutta una tecnologia ampiamente utilizzata in industria. Inoltre, con lo stesso processo si possono realizzare microelettrodi per misura o stimolazione di biopotenziali, poiché alla fine del processo di fabbricazione può essere effettuata una deposizione di materiale conduttivo (come il nitruro di titanio).
ABSTRACT
This thesis deals with design, fabrication and characterization of novel devices for minimally invasive transdermal applications, for both withdrawing -testing liquid and improving transdermal drug delivery.
At first, a new concept glucose self-monitoring device, for measuring glycaemia without causing pain or bleeding is presented. This device, called glucose pen, allows point-of-care glucose monitoring in a minimally invasive way, by avoiding both bleeding and pain caused and reducing, at the same time, the risk of infection in respect to the conventional method of glucose self monitoring. The glucose pen contains a silicon dioxide microneedle array, which samples interstitial fluid after being inserted into the skin, and a glucose biosensor for testing the sampled drop. Both fabrication process and mechanical characterization of the microneedle array is reported. Besides, the design of all the glucose pen components (circuit for glucose related signal driving and read-out, glucose sensor and microneedle array assembling into a disposable head, system for clamping the disposable head to the pen, mechanical actuating system for microneedle movement and insertion into the skin) is showed.
Then, a fabrication process based on the silicon electrochemical micromachining technology for realizing out-of-plane microneedles is reported. The developed process allowed to fabricate both sharp and flat microneedle array, to be inserted into the skin so as to improve skin permeability to a model drug and, as a consequence, the amount of drug delivered in a transdermal path.
Finally, the design and fabrication of a minimally invasive device for transdermal injection and sampling applications, carried out at Dimes Technology Center, Delft (NL), is reported. The proposed devices are fabricated by means of a wafer level process based on a Deep Reactive Ion Etching and consist in an array of hollow flat silicon microneedles, connected with a reservoir, integrated on the back-side of the device itself. The realized devices, which are very different in microneedle shape, inner diameter and array density, can be used for both drug delivery or for
in-situ fluid analysis. Besides, the proposed process is suitable to perform an
industrial production of microneedle array, since it is wafer level and is based on a well known industrial process. A further application of these devices could be sensing or stimulating biopotentials, since they can be covered with conductive layer, as titanium nitride.
CONTENTS
SOMMARIO ... I
ABSTRACT ... III
CONTENTS ... V
INTRODUCTION ... 1
1. MICRONEEDLE ARRAY FOR TRANSDERMAL APPLICATIONS: A
REVIEW OF THE TECHNOLOGIES FOR DIFFERENT MATERIALS ... 3
Abstract ... 3
1.1. Introduction ... 3
1.2. Microneedle safety: reduction of pain, irritation and
microbial penetration ... 5
1.3. Classification of different microneedle structures ... 6
1.4. Approaches of microneedle-mediated drug delivery ... 6
1.4.1 Coat and poke ... 8
1.4.2 Poke with patch ... 9
1.4.3 Dip and scrape ... 9
1.4.4 Poke and release ... 9
1.4.5 Poke and flow ... 9
1.5. Microneedle fabrication processes ... 10
1.5.1 Silicon... 10
1.5.2 Silicon dioxide ... 16
1.5.3 Metal ... 18
1.5.4 Polymer ... 22
1.5.5 Others ... 27
1.6. Microneedle transdermal applications ... 31
1.6.1. Drug delivery for transdermal immunization ... 32
1.6.2. Drug delivery for therapy ... 35
1.6.3. Sensing and fluid sampling ... 39
1.6.4 Measuring and stimulating biopotentials ... 41
1.7 Conclusions ... 42
References ... 42
2.
MINIMALLY
INVASIVE
MICROCHIP
FOR
TRANSDERMAL
INJECTION AND SAMPLING APPLICATIONS ... 59
Abstract ... 59
2.1 Introduction ... 59
2.2.1 Microneedle fabrication ... 61
2.2.2 Reservoir fabrication ... 66
2.3 Microchip characterization and discussion... 68
2.3.1 Insertion tests in skin-like polymers ... 68
2.4 Conclusions ... 73
References ... 73
3. A MINIMALLY INVASIVE DEVICE FOR INTERSTITIAL FLUID
GLUCOSE MONITORING ... 77
Abstract ... 77
3.1 Introduction ... 77
3.2 Disposable head components design ... 78
3.2.1 Microneedle array design and fabrication ... 78
3.3.2 Glucose biosensor design, fabrication and test... 79
3.2.3 Disposable head performances ... 81
3.3 Electronic circuit for the glucose-dependent signal
processing ... 82
3.3.1. Design and implementation of the electronic circuit ... 84
3.3.3 Implementation of low power mode ... 88
3.4 Sensor testing by means of the electronic circuit ... 88
3.5 Electronic circuit implementation on an printed circuit
board ... 90
3.5.1 PCB design and fabrication ... 90
3.6 Design of the disposable head and the glucose pen ... 92
3.6.1 Disposable head design ... 93
3.6.2 Glucose pen design ... 94
3.7 Glucose pen and disposable head fabrication ... 97
3.8 Conclusions ... 99
References ... 100
4. A VERSATILE ROUTE FOR THE FABRICATION OF SILICON
MICRONEEDLE ARRAY BY ELECTROCHEMICAL MICROMACHINING
TECHNOLOGY FOR DRUG DELIVERY APPLICATIONS ... 101
Abstract ... 101
4.1. Introduction ... 101
4.2. Microneedle design and fabrication ... 102
4.2.1 Fabrication process ... 103
4.2.2 Fabrication process results ... 104
4.4. Insertion tests in polymeric material simulating tissues 107
4.5. Transdermal permeation improvement by microneedle
insertion ... 111
4.5.1 In vitro permeation tests of caffeine... 112
4.5.2 Analysis of permeation test data ... 112
4.5.3 Permeation test results ... 113
4.6. Conclusions ... 114
4.7 Material and Methods ... 115
4.7.1 Design rules ... 115
4.7.2 Starting material ... 116
4.7.3 Pattern definition ... 116
4.7.4 Backside Illumination Electrochemical Etching ... 117
4.7.5 Isotropic phase to remove sacrificial structures ... 118
4.7.6 KOH etching to remove sacrificial structures ... 119
4.7.7 Theoretical analysis ... 119
4.7.8 Insertion tests setup ... 120
4.7.8 Polymeric specimen preparation ... 121
4.7.9 Skin permeation tests ... 121
References ... 121
5. FABRICATION OF A MINIMALLY INVASIVE DEVICE FOR
TRANSDERMAL INJECTION AND SAMPLING APPLICATIONS BY
MEANS OF A DEEP REACTIVE ION ETCHING- BASED PROCESS ... 124
Abstract ... 124
5.1 Introduction ... 124
5.2 Microchip design and fabrication ... 125
5.2.1 Masks design ... 127
5.2.2 Fabrication process ... 129
5.3 Experimental results ... 133
5.4 Conclusions ... 136
References ... 137
CONCLUSIONS ... 139
INTRODUCTION
Recently, microneedle arrays have been widely investigated as minimally invasive devices that would overcome the disadvantages of using hypodermic needles, for both delivering drugs into deeper skin layers and withdrawing biological fluids. The first advantages of replacing hypodermic needles with microneedles are the lack of pain and the reduced possibilities of infections, thanks to microneedle small dimensions (size from tens to hundreds of microns and length from hundreds of microns to millimeters). Particularly, microneedles became the subject of significant research starting from the mid-1990's, when microfabrication technology enabled their manufacture, with the same fabrication processes used for realizing integrated circuit.
At first, a great interest was shown at developing many fabrication processes for realizing microneedles that are very different in shape (sharp, tipped, flat), dimensions, structure (solid or hollow) and material, as silicon, silicon dioxide, metals, polymers, sugar, etc. At the same time, the advantages of microneedle applications have been evaluated and effectiveness, reliability and safety of microneedle based devices have been demonstrated.
Then, microneedle-based devices have been proposed for different applications: mainly, microneedles have been used for delivering a wide range of different low molecular weight drugs and biotherapeutics (as insulin, hormones and DNA). Particularly, among the drug delivery applications, microneedles-mediated influenza vaccination is spreading in clinical use, but also microneedle-based products were proposed and sold for cosmetic purposes.
Moreover, microneedles have been proposed for drawing fluid or performing in situ analysis, as well as measuring and stimulating biopotentials. Among the sensing applications, a great interest has been shown in glucose measurement in a pain-free way, by analyzing the glucose content in situ or in a drop of fluid sampled by microneedles: this kind of applications is able to improve diabetic patient compliance to the self-monitoring and, as a consequence, to increase the adherence of glucose monitoring frequency with the recommended levels (3-4 glucose measurements per day). By improving the glucose monitoring regimen, hypoglycaemia can be detected or prevented and the risk of long-term complications can be reduced.
Finally, by implementing both sensing and delivering operations into the same chip, containing also an integrated circuit, a closed loop system for in situ analysis and therapeutic molecule administration could be realized, going further Lab on a Chip devices. An example of a closed loop system on a chip is the pancreas on a chip, i. e. an integrated circuit that, by means of a microneedle array, delivers a precise amount of insulin after measuring in situ a certain value of glucose concentration. As a consequence, microneedle-based chips to be used for both analysis and delivering, and to be fabricated with the same process of an integrated circuit have been also investigated.
In this thesis, novel devices for minimally invasive transdermal applications are presented, for both liquid withdrawing and testing, and improving transdermal drug delivery. The proposed devices, fabricated by means of typical processes of MEMS field, are typically in silicon or silicon dioxide and have been tested in vitro.
In chapter 1 a review of state-of-the-art devices, with particular reference to fabrication processes for different proposed materials, is shown; chapter 1 summarizes also the typical applications and main results obtained in literature for microneedle-based devices.
Chapter 2 reports design, fabrication process based on electrochemical micromachining technology, and mechanical characterization of minimally invasive microchips to be used for transdermal injection/sampling applications: the chips consist in an array of silicon-dioxide hollow microneedles, with lateral size of a few micrometers, placed on the front-side of a silicon die, sticking out from the silicon surface for a hundred microns, and in connection with a reservoir grooved on the back-side of the same die.
Chapter 3 describes a first application of the microneedle array proposed in chapter 2. Particularly, the proposed device is a new-concept microneedles-based glucose pen, which allows point-of-care glucose monitoring by minimally invasive and painless operation by avoiding the invasive spot blood glucose measurement. The glucose pen contains a silicon dioxide microneedle array for sampling interstitial fluid and a glucose biosensor for testing the sampled drop. Besides, the glucose pen contains an electronic circuit for elaborating the glucose sensor signal and displaying the glycaemia value.
In chapter 4, a versatile route for the fabrication of arrays of out-of-plane silicon microneedles is described. The fabrication process is based on the silicon electrochemical micromachining technology and allows to realize microneedles, with height of about hundred microns and microneedle centre-by-centre distance ranging from 50 to 400 µm. The proposed devices are characterized by in-vitro insertion tests, both on skin-like polymers and hairless-mouse skin, and permeation tests, which demonstrate that microneedle treatment results in an improved transdermal permeability to a model drug.
In chapter 5 the results of a research activity carried out at Dimes Technology Center, Delft (NL), are reported. The design and fabrication of a minimally invasive device for transdermal injection and sampling applications are described. The devices are fabricated by means of a wafer level process based on a Deep Reactive Ion Etching and consist in an array of hollow flat silicon microneedles, connected with a reservoir, integrated on the back-side of the device itself. The realized devices are very different in microneedle shapes (cylindrical or tipped), inner diameter (from 10 to 50 µm) and array density (from 625 to 40000 needle\cm2) and can be used for both drug delivery or fluid sampling. Finally, conclusions summarize the thesis and show perspectives for further research.
1.
MICRONEEDLE
ARRAY
FOR
TRANSDERMAL
APPLICATIONS: A REVIEW OF THE TECHNOLOGIES FOR
DIFFERENT MATERIALS
Abstract
Recently, microneedles have been widely investigated as a minimally invasive device to both delivery drugs in deeper skin layers and withdraw biological fluids, without causing pain. Particularly, different fabrication processes of microneedles have been investigated, in order to furnish different alternatives in shapes, dimensions and materials for different applications. In this chapter, after a brief discussion about advantages of microneedle-based drug delivery, a review of state-of-the-art devices and their applications will be shown.
1.1. Introduction
Microneedles became the subject of significant research starting from the mid-1990's when microfabrication technology enabled their manufacture and, as a consequence, they could be investigated as a promising devices for actuating drug delivery or fluid sampling by overcoming limitations of hypodermic needles. The first disadvantage of hypodermic needles, that can be avoided by using microneedle-based devices, is the needle phobia, which affects about 10% of adult population and in greater extend children [1, 2]. At first, microneedles are short enough to avoid encountering pain receptors, residing beneath the skin’s outer layer, so that the actual pain caused in the patient is reduced, increasing patient compliance. Besides, the small dimensions of caused wounds reduce the risks of infection, bleeding and irritation.
Particularly, microneedles for drug delivery applications are considered minimally invasive device to overcome the skin barrier by improving the limited permeability of molecules more weighty than 500 Da [3]. Microneedle employment actually performs a hybrid approach combining both transdermal patches, which are affected by reduced drug penetration, and hypodermic or intramuscular needles: actually, microneedles are able to delivery drug into deeper tissue layers, through microchannels created by microneedle insertion, by increasing the amount of drug released in respect to transdermal patches. Moreover, microneedles avoid the most important issue of the oral administration that is the drug destruction in gastro-intestinal tract because of enzymatic activity [4, 5].
The pain and irritation reduction, as well as infection avoidance, as mentioned above, are probably the most important advantages of microneedle application, in respect to hypodermic and intramuscular injection. Moreover, microneedle-based drug delivery increases the perspectives of controlled release and the possibility of patient self-administration, without needing to be trained by specialized personnel. For all these reasons, microneedle application is considered a valid alternative to conventional routes of drug delivery, as oral administration, transdermal patch and intradermal or intramuscular injections.
The growing interest on microneedle applications is clearly shown by the increasing number of works about this topic, widespread in different subjects (showed in fig. (1.1) according to results obtained from [6]). Particularly, fig. (1.1.a) clearly shows the increasing interest in this research topic since the last decade,
while fig. (1.1.b) shows that microneedles were mainly investigated in engineering area, followed by the pharmacological area. The gap between these main areas of interest is due to the procedure generally used to design a clinical trial [7].
First of all, before investigating a new drug or drug delivery system on human beings, investigations on a safe and reproducible fabrication process for microneedle realization have to be carried out. Afterwards, some useful preclinical information about nonhuman efficacy, toxicity and pharmacokinetic are generally obtained from in vitro and in vivo animal models. Finally, an analysis of randomized
Fig. 1.1 – Number of publications about microneedles divided by year of publication and subject area [6].
0 20 40 60 80 100 120 140 160 1952 1972 1975 1977 1980 1982 1985 1987 1989 1991 1993 1995 1997 1999 2001 2003 2005 2007 2009 2011 2013 YEAR
trials is performed in order to confirm the effectiveness of new protocols. The results of this analysis, generally consisting in several phases of clinical research, are then published during the path to the final approval of the device, which in some cases can allow the commercialization of the products (section 1.6.2.4).
1.2. Microneedle safety: reduction of pain, irritation and
microbial penetration
The main advantage of microneedle application is pain reduction, which depends on microneedle dimensions, kind of application and biocompatibility of the material microneedles are made out. Particularly, the pain caused was investigated in relationship with microneedle dimensions and number or infusion pressure [8, 9, 10, 11, 12, 13]. Gill et al. compared pain induced by a 26-gauge hypodermic needle and different microneedle arrays (steel microneedle array with length ranging from 480 to 1450 µm) [12]. Pain scores, corresponding to all the needles tested, were compared resulting that the microneedle pain scores are in the range from 5 to 40% of the hypodermic needle score. Therefore, microneedles are actually less painful than the hypodermic needle, even if the microneedle number influences pain scores. Similar results were shown by Gupta et al. by comparing the pain caused by glass hollow microneedles (length of 500 µm, 750 µm, 1 mm and 4 mm) [9]. The pain induced is also influenced by infusion pressure. Particularly, the pain caused by a low flow rate is significantly less than that caused by high flow rate. Besides, pain at the low flow rate is actually independent by drug volume infused, therefore pain can be kept low by reducing infusion pressure, even when a great drug volume is delivered.
Other approaches to evaluate safety of microneedles are the evaluation of redness and blood flow caused or dimensions of microchannels created [14, 15, 16, 17]. All the studies concluded that microneedle-mediated transdermal drug delivery is safe and well tolerated; moreover, the created conduits collapse without leaving injuries. Particularly, Bal et al. evaluated the safety in terms of skin redness and blood flow of solid steel microneedles (with a length of 200, 300 or 400 µm) and microneedle arrays assembled from 30-gauge hypodermic hollow metal needles (with a length of 300 and 550 µm). Shape and length of microneedles influence the degree of irritation, which is minimal with respect to tape stripping for all the devices tested and lasts less than 2 hours [8]. Noh et al. obtained similar results about microneedle caused irritation by studying the skin redness after application of 500 µm tall polymer microneedles on human skin. Particularly, the redness caused increased during microneedle insertion and then recovered to baseline after microneedle removal. The authors noticed a little difference in redness reduction after applying microneedles for various times (from 2 minute to 4 hours), but redness was generally maintained until 30 minutes and then rapidly decreased between 30 minute and 2 hours [17]. Finally, Van Damme et al. evaluated the irritation caused by microneedle-based vaccine delivery and showed that such method of drug administration is effective, safe, and reliable [15].
As well as pain and irritation, microneedles decrease the actual microbial penetration into the tissue with the respect to hypodermic needle. Donnelly et al. investigated in vitro microbial penetration in microneedle-created conduits, by comparing the results obtained after both hypodermic injection and microneedle application [18]. Authors concluded that microneedle puncture causes a
significantly less microbial penetration than hypodermic needle puncture. Moreover, in microneedle-punctured skin no microorganisms cross the viable epidermis, in contrast to needle-punctured skin. This feature makes microneedles very attractive for future transdermal drug delivery use in clinical practice, since their application can avoid either local or systemic infection in normal conditions.
1.3. Classification of different microneedle structures
In order to properly understand the fabrication process of state-of-the-art microneedles, it is useful classified microneedle structure: particularly, microneedle arrays can be classified in in-plane and out-of-plane, while microneedle structure can be solid or hollow.
The difference between in-plane and out-of-plane microneedle array concerns the position of the needles longitudinal axis with respect to the surface of the substrate from which the needles were produced: in-plane arrays show a longitudinal axis parallel to the surface of the substrate, while for out-of-plane array the longitudinal axis is perpendicular to the substrate surface. This classification has a consequence in term of needles design and fabrication process: in-plane needles have no limitation in term of length, while, as a consequence of the fabrication process, there are strong constrains in term of transversal section shaping and resizing along the longitudinal axis. On the other hand, for out-of-plane needles the length is limited at thickness of the substrate from which it is produced, while no significant constrains are present for transversal section, with the possibility to realize complex shapes and to resize section dimensions along the longitudinal axis. Typically, a microneedle-based device for drug delivery is an out-of-plane array that shows the needles arranged in a two-dimensional array structure. For this reason, fabrication process of microneedle-based devices starting from an in-plane array requires post-processing step. The fabrication process, in this case, can be divided into two principal steps: 1) the production of one-dimensional arrays of needle (in-plane arrays) and 2) their assembling in a two-dimensional array. This process causes a low control of the dimension of the array in the assembling direction, depending on the assembly phase itself. In-plane microneedle can be, on the other hand, used for recording electrical activity or delivery chemical substances in cells or neurons [19, 20, 21], while for transdermal application out-of-plane microneedles are generally used.
Besides, microneedles can be classified with respect to their structure: particularly, hollow microneedles are provided of an inner channel, parallel to the microneedle longitudinal axis that is not present in solid microneedles. Generally, the channel connects microneedle tip with a reservoir placed in the backside of the substrate and often its fabrication process is more complex than that required for solid microneedles. Special kind of solid microneedles are dissolving: actually, they are solid microneedles fabricated with some special materials, able to dissolve itself after being inserted into the skin.
1.4. Approaches of microneedle-mediated drug delivery
Several approaches can be employed to use microneedles for drug delivery (Fig. 1.2), according also with microneedle solid, hollow or dissolving structure [4, 5]. In the case of solid or hollow microneedles without connection with a reservoir, the techniques used are “coat and poke”, “poke with patch”, “dip and scrape”. For
dissolving or hollow microneedles "poke and release" or "poke and flow" technique are respectively used. All these techniques are schematically shown in fig. 1.2.
Fig. 1.2 – Different methods for delivering drug by means of microneedles (a-c): (a) coat and poke, (b) poke with patch, (c) dip and scrape.
1.4.1 Coat and poke
“Coat and poke” (fig. 1.2.a) is the most common technique and consists in inserting solid drug-coated microneedles. This approach requires optimization of drug formulation to form stable, uniform and thick coatings and to avoid drug degradation [22, 23, 24, 25, 26, 27, 28]. The actual coated drug volume depends on the liquid wetting behaviour on the needle surface, the formulation viscosity, the surface tension and the method of coating. By applying this approach, a small drug volume can be released, because all the drug deliverable has to be coated on the microneedle surface. The available volume can be increased by creating pokes or holes into microneedles [29]: in this way, an array exploits pocket of different shape and size that can be selectively filled with different drugs. Moreover, different pocket depths result in different drug delivery depths and allow a precise localization of the release site. Generally, “coat and poke” approach is used to deliver in transdermal route vaccines or genes to exploit immunogenic skin response [28,30, 31, 32, 33, 34, 35]. Some authors applied this approach to deliver other drugs [36, 37, 38]. Regarding vaccine, long-term stability of microneedles coated with inactivated influenza vaccine was studied by Choi et al. [39]. Authors explained that, over a time period of many weeks, crystallization and phase separation of the vaccine coating damages the vaccine and can render it completely non-immunogenic. For this reason, they proposed that inhibition of the phase changes in the vaccine coating film is necessary for the development of
Fig. 1.2 – Different methods for delivering drug by means of microneedles (d-e): (d) poke and release, (e) poke and flow.
long-term stable vaccine-coated microneedles, while for freshly prepared microneedles this issue is avoided.
1.4.2 Poke with patch
“Poke with patch” technique (fig. 1.2.b) consists of two steps: at first microneedles are inserted and removed from the tissue, then microchannels created by microneedles are exploited to improve drug diffusion from a transdermal patch, placed on the tissue surface, to tissue internal layers. This approach is used for delivering drugs or vaccine and is generally carried out through solid microneedles [40, 41, 42, 43, 44, 45, 46, 47, 48].
1.4.3 Dip and scrape
“Dip and scrape” technique (fig. 1.2.c) consists in dipping microneedles in a drug solution, then scraping them on tissue surface. Scraping creates microabrasions through which drug can diffuse more efficiently than in untreated skin [49]. This approach has been proposed to deliver genes, but is less common than previous methods.
1.4.4 Poke and release
“Poke and release” (fig. 1.2.d) approach is completely different and has been recently investigated [50, 51, 52, 53, 54, 55, 56, 57, 58, 59, 60, 61, 62]. It consists in encapsulating drug into dissolving microneedles, so that both microneedles and drug dissolve when inserted into the target tissue. In such a way, drug delivery can be controlled varying the dissolving rate of the microneedles and the available drug volume can be increased by increasing microneedle dimensions or number. Recently, a different approach for increasing drug delivered by dissolving microneedles was proposed by Ito et al. [63]. They proposed to load the drug alternatively in the whole dissolving microneedles, or in both dissolving microneedles and chip. The latter alternative system was demonstrated to be useful for the delivery of more than 1.0 mg of the drug to the systemic circulation. Dissolving microneedles are usually fabricated from biodegradable polymers or polysaccharides and allow biocompatibility to be improved. Besides, safety problems can be solved since microneedles dissolve in few minutes into the skin, without leaving sharp tips that can cause damage.
1.4.5 Poke and flow
“Poke and flow” (fig. 1.2.e) is feasible only by using hollow microneedles. It consists in applying a hollow microneedle array that is in communication with a drug filled reservoir. In this way, drug can flow from the reservoir to deeper tissue layers, by means of either diffusion or pumping systems [64, 65, 66, 67]. The pumping system can increase the rate of drug delivery [68] or allow the controlled release [65, 69]. In the last case, the pumping system has to be externally controlled to realize a closed-loop therapy.
The same microneedle structure needed for “poke and flow” technique allows to perform also fluid drawing, as explained in section 1.6.2; for this kind of application, thanks to smaller size of hollow microneedle channels, the drawing can be performed without an external pump, but simply by means of capillarity effect [70].
1.5. Microneedle fabrication processes
The choice of the microneedle material for transdermal applications should be based on criteria such as biocompatibility, mechanical strength for insertion into the skin and compatibility with a reproducible high-yield fabrication process for low cost production. The first material proposed for microneedle production was silicon: this choice allows physical strength and fabrication processes directly resulting from the well-established integrated circuit technology, even if silicon is relatively expensive, fragile and unproven as biocompatible material. Nevertheless, the interest for this material in the biomedical field is always high for the possibility offered by silicon technology to fabricate micrometric devices by using other biocompatible materials. As a matter of fact, the majority of the microsystems produced with biocompatible materials (like polymers) uses silicon structures as masters. Besides, silicon dioxide microneedles are proposed to overcome the biocompatibility limitation of silicon while keeping the same features and exploit similar fabrication processes. A further approach is the fabrication of metallic microneedles, but they are generally expensive, non-biocompatible and brittle. Polymer microneedles overcome the limitations of silicon and metal microneedles providing advantages like a promising low cost fabrication process, mechanical strength, improved resistance to shear-induced breakage due to polymer viscoelasticity and safety in case of accidental breakage of needle in the skin.
1.5.1 Silicon
Initially, the material selected for microneedles fabrication was silicon because the first fabrication processes resulted from micro-electro-mechanical systems (MEMS) technology [19, 20, 44, 69, 70, 71, 72, 21, 73, 74, 75, 74, 76, 77, 78, 79, 80, 81, 82, 83, 84, 85, 86, 87, 88, 89, 90, 91, 92, 93, 94, 95, 96, 97, 98, 99, 100, 101, 102]. Generally, the processes for silicon microneedles fabrication are based on three basic technological steps: 1) deposition/growth of a protecting layer, 2) patterning the protecting layer and 3) chemical etching of silicon substrate [103]. The deposition/growth step refers to the production of a thin film of several materials on the substrate: this film can be a metal (i.e. aluminium, chromium, gold etc.), an insulator (i.e. silicon dioxide, silicon nitride etc.) or soft materials (i.e. different kind of polymers), with a thickness ranging from a few nanometers to tens micrometers. Each of these different materials has a specific deposition/growth technique: typically, metallic films are deposited by means of an evaporation step under vacuum condition, while the insulator layers can be realized with a chemical vapor deposition step or, for the silicon dioxide, with a thermal oxidation. Finally, soft materials are generally spun on the substrate. The patterning step is used to define a design onto the protecting layer: when the protective layer is made out a soft material the patterning is performed by means of photolithography, electron beam lithography, ion beam lithography or X-ray lithography [103]. The most common and flexible lithographic step is the photolithography: in each case, this step is used to transfer a pattern into the soft material layer by selective exposure to a proper source (radiation, electron, ion or X-ray). The exposure is followed by a development step, which actually creates the design on the soft material layer. On the other hand, the pattern definition on a hard material protecting layer requires the presence of a soft material layer onto the hard material layer. In this way, the pattern can be defined previously on the soft layer, which plays the role of a mask
layer: trough this mask layer, the pattern is transferred in the protective layer by means of specific etching of the hard material. After this phase, the etching step is performed: it consists in a dissolution process that allows the realization of a design into the substrate by removing material from the substrate itself through the unprotected parts of its surface [103]. Wet or dry etchings are the two possible etching ways: wet etching is performed in a liquid phase, while dry etching is performed in gaseous media.
Starting from a silicon substrate, by properly combining the above mentioned technological steps, it is possible to fabricate silicon in-plane or out-of-plane, solid or hollow microneedles: fig. (1.3) schematically shows the process flow required for having in-plane, out-of-plane, solid and hollow microneedles.
1.5.1.1 Silicon array of in-plane microneedles
At present, two are the processes proposed for the production of silicon in-plane solid/hollow microneedles: the boron etch-stop technique [19, 20, 21, 72, 73] and a dry etching based method [74, 75, 97].
Proposed for the first time by K. Najafi et al. [19, 20] in the 1985 for the fabrication of silicon solid in-plane microelectrodes, the boron etch-stop technique is the result of a selective doping of a silicon substrate for the needles structure definition,
Fig. 1.3 – Main steps of silicon microneedle fabrication process: (a) flowchart for in-plane microneedle fabrication; (b) flowchart for out-of-plane solid microneedle
followed by their release from the substrate, by means of a silicon wet etching (fig. (1.3.a)). The first step of this fabrication process consists in depositing a hard mask layer (mainly silicon dioxide) on the substrate surface and patterning it in order to transfer the needles geometry into the mask layer, which in this case corresponds to the protecting layer. The following step is boron doping of the wafer through the windows opened into the mask layer: in this way, the complete microneedle structures can be defined into the substrate as a heavily p+ doped region in the silicon wafer. Particularly, by properly tailoring the doping process parameters, microneedle thickness can be set. Generally, the doping process causes a change electrical properties of silicon substrate by introducing in the lattice specific species, i.e. phosphorous or boron, that allow silicon becomes more conductive in term of electron, n-type silicon, or hole, p-type silicon, respectively. The doping process, by means of a thermal diffusion of the dopant species implanted in the substrate or deposited on its surface, changes the substrate doping only in a region close to the surface for a thickness ranging from a few micrometers to tens micrometers [103]. In the process proposed, the doped region is the microneedle structures to be realized. After the mask layer removing, a silicon wet etching in ethylene diamine-pyrocatechol (EDP) is performed on silicon substrate [71]. The selectivity of EDP etch between the silicon substrate and the heavily p+ doped regions [104] allows the wafer dissolution and the release of the microneedle structures. Fig. (1.4.a) shows a SEM image of the tip of the in-plane silicon microneedle presented by the K. Najafi group. The proposed microneedles have a length ranging from 1.5 mm to 3 mm and a cross section of about 50 µm x 15 µm at the tip.
By adding few steps of deposition, patterning and silicon etching before the EDP etch, it is possible to define metallic terminals on the microneedles, so as to enable their application as electrode recording [19, 20, 72], or to convert the solid microneedles into hollow microneedles [21, 41]. As an example, the in-plane hollow needles proposed by Lin et al. [6] were ranging from 1 mm to 6 mm in height and 80 µm in width, while the channels were ranging from 30 to 50 µm in width and 9 µm in height.
On the other hand, Izumi et al. proposed a simpler and more versatile process for the fabrication of silicon solid in-plane microneedles with a three-dimensional sharp tip. This process replaces EDP etch, stopped by a boron doping, with a dry etching in a fluorocarbon atmosphere [74, 97]. As the previous technique, the process initiates with deposition and patterning of a hard mask layer (aluminium [97] or silicon dioxide [74]), but in this process the mask layout is complementary to that of shown in fig. (1.3.a) (red layer). Actually, in this case the region corresponding to microneedles, defined by the mask layer patterning, have to be protected from silicon etching performed in the following step. Then, a silicon anisotropic dry etching, named deep reactive ion etching (DRIE) allows the needles structure to be directly realized, by skipping the boron doping step (fig. (1.3.a)) [103]. By properly selecting the gas composition and the parameters of the dry etching, it is possible to drive the silicon removing process mainly along the direction perpendicular to the substrate surface, avoiding the dissolution under the mask layer. Finally, the microneedles are released from the substrate with a silicon wet or dry etching from the substrate back side. This process overcomes the boron etch-stop technique limitation in term of needle cross section height and shape. While, even in this case, the height is limited only by the substrate thickness, this process is provided
of a good flexibility in the cross section shaping by means of tuning the DRIE anisotropy and selecting a proper mask pattern. As an example, Izumi et al. [97] produced in-plane solid microneedles with a jagged or harpoon shape more than 1 mm length and less than 100 µm wide.
1.5.1.2 Silicon array of out-of-plane solid microneedles
Fig. (1.3.b) shows schematically the main technological steps of the typical process for the production of a two-dimensional array of silicon solid out-of-plane microneedles [16, 44, 81, 82, 83, 85, 96, 98, 99]. The process can be synthesized in a selective anisotropic removing of silicon from a wafer trough a proper mask layer. The first step of the fabrication process is the deposition and patterning of a hard mask layer (i.e. silicon dioxide, silicon nitride or a metal) onto the silicon substrate surface. The mask layout is a dot array defining the relative position in the array and the transversal section of the needles. The following anisotropic wet or dry etching removes the silicon from the unprotected surface regions, deeply and slightly in the perpendicular and parallel direction with respect to the substrate surface, respectively. The height of the resulting needles is a function of the removing process depth, while the shape and the dimensions of the transversal section along the needle axis depend on mask and removing process anisotropy. The resulting solid needles have the longitudinal axis perpendicular to the substrate surface and are anchored to it at the bottom.
S. Henry et al. [99] proposed, for the first time with the described process, a 20 by 20 and 150 µm spaced array of silicon out-of-plane solid microneedles (fig. (1.4.b)). By using an anisotropic DRIE, the authors obtained 150 µm tall conical microneedles with an extremely sharp tips, whose radius of curvature is lower than 1 µm. On the other hand, by alternating isotropic and anisotropic etching, P. Griss
et al. [98] demonstrated the possibility to produce cylindrical microneedles, with
height and diameter ranging from 100 to 200 µm and 30 to 50 µm, respectively, that are provided of a sharp tip radius less than 0.5 µm. Particularly, this process exploited the anisotropy of DRIE for producing the cylindrical structure of needles, while the isotropic etch phase was used to shape the needle tip for making it sharp. A similar process, with a different order of wet and dry etching steps, was proposed by Li et al. for the fabrication of super-short microneedles with a length of 70– 80 µm, by means of an isotropic wet etching and a DRIE [44].
Wilke et al. [85] reported two different processes for silicon out-of-plane solid microneedles fabrication: they obtained microneedles tall 300 µm by using an anisotropic wet etching or different steps of isotropic and anisotropic DRIE etching (the latter process is similar to that reported by Griss et al.). Differently from the dry etching (as DRIE), the anisotropy of wet etching is due to the crystallographic orientation of silicon [103], therefore the resulting needles were pyramidal, instead of being cylindrical. Besides, the array pitch, that is the centre-to-centre distance between two adjacent needles, could not be lower than the needle height.
A maskless process for the production of out-of-plane solid microneedles by wet etching was proposed by Shikida et al. [79, 80]. In this case, the pattern definition was made up by using a dicing saw technique. After the formation of a two-dimensional array of prismatic pillar on the silicon substrate, by using a diamond blade, an anisotropic wet etching was carried out in order to convert the pillars into pen-shaped microneedles with an height of 300 µm, a pitch of 100 µm and a sharp tip with a radius less than 0.1 µm. Although straightforward, the proposed process had strong limitation in pattern geometry, needle pitch and dimensions, due to the
limitation of the dicing saw technique. A similar process was proposed by Yan et al. [78] for fabrication of microneedle array witch a length of 700 µm, a pitch of 600 µm, and a size of 400 µm and 30 µm, in the bottom-side and in the top-side respectively. Particularly, authors proposed to produce tapered pillars, trough two different beveled blades designed with a bevel angle of 20°; afterwards, the cut microneedles were placed in a mixed solution (HNO3:HF = 19:2 in volume) to
sharpen tips and smooth microneedle surface.
An interesting solution to improve drug delivery by silicon solid microneedles was proposed by Ji et al. [82]. The out-of-plane needles shown were fabricated by means of a dry etching and provided of a porous tip, where a drug could be loaded. The porous silicon tip was produced after the needle array definition, by masking the body needle with a proper layer (silicon nitride) and performing an electrochemical etching of the uncovered tip. This solution increased the volume of the drug to be delivered with respect to a simple needle coating; moreover, this structure showed also the advantages of porous silicon as the low toxicity and degradation properties that make it more biocompatible than silicon [105]. The pyramidal silicon microneedles proposed were 100 to 150 μm in height, while their porous tips were 30 μm tall.
Solid out-of-plane microneedles can be realized by performing an HNA (hydrofluoric acid–nitric acid–acetic acid) etching of silicon, instead of the other etching above mentioned [95]. By means of this approach, Hamzah et al. realized solid silicon microneedles with an average height of 159.4 μm, an average base width of 110.9 μm and a tip angle and diameter of 19.2°and 0.38 μm, respectively. Finally, Islam et al. [94] proposed a very different fabrication approach for realizing out-of-plane solid microneedles: these needles were grown by means of vapor-liquid-solid (VLS) method, starting from a boron doped (111) oriented silicon. In this case, at first a pattern of gold dots, used as metallic catalyst, was deposited and patterned in specific sites of silicon substrate. Afterwards, a VLS grown of silicon was carried out into a vacuum chamber, where the silicon substrate was heated, so as to form gold-silicon liquid alloy droplets on the silicon surface, thanks to the mixture of gold and silicon atoms of the substrate. By adding proper gases (Si2H6)
into the chamber, silicon atoms from the gas were absorbed by the gold-silicon liquid alloy droplet, which acted like a trap for the gas. When the droplet became supersaturated, silicon atoms began to precipitate at the interface of alloy droplet and substrate and hence needle-like silicon crystals were grown perpendicular to silicon surface.
1.5.1.3 Silicon array of out-of-plane hollow microneedles
In order to convert silicon microneedles from solid to hollow, additional technological steps are required to the above mentioned process, fig. (1.3.c). Hollow needles are obtained by defining micrometric channels along the direction perpendicular to the substrate surface connecting the needles tip with the substrate backside. For this purpose, the hollow out-of plane microneedles fabrication process involves two mask layer: one for the microneedles definition and one for the microchannel extrusion; generally the former is defined onto the substrate frontside, while the latter is defined onto the substrate backside. The alignment between the two mask layers represents a critical point due to the micrometric dimension of the needles: a slight alignment error could cause that needles do not work properly. By introducing further deposition and patterning steps of a hard
mask and silicon etching steps to the above mentioned solid out-of plane microneedles processes, hollow needles can be produced. In order groove the channels into the substrate, a silicon anisotropic DRIE is generally used, while for microneedle structure definition can be used both wet [89] or dry etching of silicon [69, 100]. The DRIE step is carried out through the openings into the mask layout, which is generally a matrix of hole defining the microchannels width and position. This kind of process can also allow, without any additional steps, to sharp microneedle tips by changing the selectivity between silicon and mask layer, during the etching of the microneedle shafts. This capability was demonstrated by Khanna
et al. [101, 102] that fabricated out-of-plane hollow microneedles by means of two
DRIE steps and sharpened microneedle tips in different way by varying DRIE recipe. Particularly, they fabricated microneedles with internal and external diameter ranging from 85 μm to 119.8 μm and from 118.8 μm to 160 μm, respectively, with a length of about 120 μm.
Mukerjee et al. [70] presented silicon out-of-plane hollow microneedles produced by combining DRIE, diamond blade circular sawing and isotropic etching. Particularly, the process consists in two anisotropic DRIE steps, one from the substrate backside, for channel etching, and one from the substrate frontside for the microneedle structure definition. Particularly, the frontside wet etching was performed after patterning the substrate surface by means a dicing saw; this etching step defines the needles structure and sharpened microneedle tips. By shifting the backside mask respect to frontside mask, different shapes of needles were produced: particularly, the volcano-like design and micro-hypodermic design were both 200 to 350 µm high, 120 µm wide at the base and 300 µm in pitch, with a channel diameter ranging from 10 to 15 µm, but they differed for position of the hole, which was centred in the former and shifted of 25 µm from the column’s centre in the latter. For this reason, micro-hypodermic design showed an extremely sharp tip (2 µm radius of curvature) that could not be fabricated in the volcano-like design.
A different approach was presented by Gardeniers et al. [87] for fabricating complex-shaped hollow microneedles in silicon. The process, consisting of a sequence of DRIE, anisotropic wet etching and conformal thin film deposition steps, useful for fabrication of a large variety of needle shapes by means of a three-dimensional mask. Particularly, at first the needles structure was defined by means of anisotropic DRIE etching, afterwards by means of a layer deposition step (silicon nitride) a three-dimensional protective mask was created through needles channel, finally by using an anisotropic wet etching the final definition of the structure was performed. This approach overcame limitation of wet etching with proper three-dimensional mask and allowed to fabricate, for example, needles 350 µm high, with a triangular tip shape, a base of 250 µm, and a maximum hole width of 70 µm.
P. Griss et al. [86] presented the first silicon out-of-plane hollow microneedles with openings along the shaft rather than a single hole at the tip (fig. (1.4.c)). These structures are less prone to clogging than tip opened ones during the skin penetration and the large area of the side openings allow large area of liquid exposure to the skin. In order to obtain side opened needles, the anisotropic DRIE from the substrate back side was stopped before reaching the substrate front side and then a thermal oxidation step was used to convert the channels into silicon dioxide microtubes. By means of a front side mask layer with a cross shaped array
pattern and anisotropic DRIE the needles structure was defined so that uncovering partially the silicon dioxide microtube. Finally, a silicon dioxide etching was used to remove the microtube creating opening along the needles side walls, as shown in fig. (1.4.c). These microneedles were 210 µm high, with a channel having a diameter of about 50 µm. In the same way, by changing the etching time and the mask ultrasharp microneedle, with a cylindrical shaft opened laterally, can be realized [92]. Particularly, after the etching process, the tips of the needles were sharpened by means of a wet oxidation (2 μm thick) followed by a consecutive oxide removal, allowing the realization of ultrasharp apex with a tip radius below 100 nm.
1.5.2 Silicon dioxide
In order to overcome the silicon low biocompatibility, silicon dioxide (SiO2)
microneedles were proposed [106, 107, 108, 109, 110, 111, 112, 113, 114, 115, 116]. At present, only hollow out-of-plane needles are known, probably due to their fabrication processes involving the same technological steps used for silicon needles realization. The proposed approaches, generally, use a micromachined silicon substrate as a scaffold for the needles structure definition. By partially converting the silicon of the scaffold into SiO2 and etching of the silicon between
the resulting SiO2 structures, hollow microneedles are produced. Fig. (1.5) shows
schematically the main technological steps of this process. The first step is the production of the silicon scaffold by means of the deposition/growth and patterning of a hard mask layer followed by a silicon anisotropic etching step. For out-of-plane hollow needles purpose, the etched structure is typically an array of micrometric channel starting from the front side of the substrate and perpendicular to its surface. The channels length could be as long as the substrate thickness (thorough holes) or less (one-end-closed holes). Afterwards, a silicon thermal oxidation step is used to convert the microchannels into silicon dioxide micrometric tubes. This process of oxide growth determines the sidewall thickness of the final needles: as long is the thermal step as thick is the resulting needles. Finally, by an unmasked silicon isotropic etching from the substrate front side, SiO2 out-of-plane hollow
needles protruding from the substrate are realized. Only for the one-end-closed tubes is required a second isotropic etching from the substrate backside to remove the silicon below the microtubes and an SiO2 etching to have hollow needles.
Fig. 1.4 – Silicon microneedles. Images were reproduced from (a) [20] © 1985 IEEE, (b) [99] (with permission from ref. [99]), (c) [86] © 2001 IEEE.
T. Shibata et al. [110, 111] presented an array of hollow microneedles (fig. (1.6.a)) by using an anisotropic DRIE for producing thorough holes into the silicon substrate and an isotropic etching for leaving SiO2 microneedle structures. The
needles were cylindrical with inner diameter of 3.5 µm, outer diameter of 5.5 µm, pitch ranging from 13 µm to 26 µm and length ranging from 40 µm to 70 µm. The critical step of the above mentioned process is the silicon scaffold definition. As the etched channels into the silicon substrate form the silicon dioxide needles geometry, an etching step allowing the control of the channels length as well as their cross section shape and dimensions is required. For this purpose, as an alternative to the anisotropic DRIE etching, a back-side illumination electrochemical etching (BIEE) of silicon was proposed [107]. Strambini et al. [107] demonstrated the feasibility of the n-type silicon BIEE technique to define in a simple way micrometric three-dimensional structures from a silicon substrate with aspect-ratio (that is the ratio between the perpendicular structure dimension and the lower parallel structure dimension with respect to the substrate surface) significantly higher than that obtained with an anisotropic DRIE. SiO2 microneedles produced
by means of BIEE for silicon scaffold fabrication are shown in fig. (1.6.b). These needles had squared cross section with inner side of 3.6 µm, outer side of 5.6 µm, pitch of 10 µm and length of about 100 µm. With a similar approach, Rajaraman et
al. [108] presented needles with squared cross section, two different inner side
(5 µm and 20 µm), two different pitches (20 µm and 100 µm, respectively), sidewall thickness of about 1 µm and height of about 60 µm. The two different dimensions were obtained by using the dependence of the pitch and inner side of the etched channels by the substrate doping level for the electrochemical etching technique. An alternative method was proposed by K. Takei et al. [114, 115, 116] for producing SiO2 microneedle by covering with silicon dioxide an array silicon
Fig. 1.5 – Main steps of silicon dioxide microneedle fabrication process: flowchart for typical structures, i. e. out-of-plane hollow microneedles.
column, which was, as a consequence, the complementary respect to the previous cases. Particularly, the silicon scaffold was an array silicon column grown from the substrate surface by means of a selective vapor–liquid–solid (VLS) technique. Differently from the other approach where the crystallographic orientation of the substrate are (100), for the applicability of VLS technique a silicon substrate (111) oriented and a pattern of gold dots on its surface are required. By means of a thermal step in a Si2H6-based atmosphere, silicon microprobes could grow in
correspondence of the gold dot, perpendicularly to the substrate surface with a cross section equal to that of the gold dot (as discussed above). Finally, by SiO2
deposition and silicon etching, microtubes with 0.5 μm in wall thickness, 20 μm height and either 2.5 μm, 4.1 μm, 4.6 μm, and 6.4 μm inner diameter were produced fig. (1.6.c). The increased complexity of such approach with respect the previous described makes it less viable.
1.5.3 Metal
Like for silicon, metallic in-plane and out-of-plane, solid and hollow were proposed [22, 91, 117, 118, 119, 120, 121, 122, 123, 124, 125, 126]. Generally, metallic microneedles are greater in dimensions than silicon microneedles, and their biocompatibility depends mostly on specific metal used, ranging from a great biocompatibility for titanium or palladium and poor biocompatibility for nickel [127].
1.5.3.1 Metallic array of in-plane microneedles
The simplest method for metallic microneedle fabrication is a laser cutting of metallic sheets (fig. (1.7.a)): particularly, in-plane solid needles can be made up by using stainless steel sheets cut by means an infrared laser with a proper energy that ablates the metal sheet creating the needles in the plane of the sheet [23, 120]. Therefore, by means of a laser-controlled beam, the specific shape and orientation of the microneedle arrays are defined. Before using microneedles, an electropolishing of the structure is necessary in order to remove debris of cutting. Afterwards, such in-plane solid needles can be converted into out-of-plane by manually bent at 90° the arrays and properly gluing them.
Fig. 1.6 – Silicon dioxide microneedles. Images were reproduced from (a) [110] © 2007 IEEE, (b) [107] - with permission from The Royal Society of Chemistry-,
Gill et al. [22] reported a patch of 50 microneedles assembled with a set of ten in-plane rows of microneedles with each row containing five microneedles 75 μm wide, laser-cut into a 75 μm thick stainless steel sheet. The microneedle length could be varied, as well as the shape, by changing the design of the laser cutting. With the same process, Martantano et al. [120] presented an array containing 7 rows of 15 needles each, for a total of 105 needles. Each needle was 50 μm by 200 μm width at the base, and tapered over a 1000 μm length to a sharp tip fig. (1.8.a).
Fig. 1.7 – Main steps of metal microneedle fabrication process: flowchart for in plane and out of plane.
For fabricating in-plane hollow microneedles, more complex processes were proposed [117, 118, 119, 124]. Generally, these processes involve a sequence of metal deposition steps and sacrificial layers etching (fig. (1.7.b)). At first, the deposition and patterning of a first metallic film are performed on a substrate (typically silicon) in order to define needle layout. Afterwards, a sacrificial layer (typically polymer) is deposited and patterned on the metallic film to form the needle inside volume and the needles side walls are completed by means a metal electroplating step. Finally, by removing the sacrificial layer with a selective etching, in-plane hollow needles resulted. Brazzle et al. [117] reported one-dimensional array of 25 palladium in-plane hollow needles with cross sectional inner dimensions of 30 μm in width and 20 μm in height, outer dimensions of 80 μm in width and 60 μm in height fig. (1.8.b). Each needle showed 2 mm long channel, a wall thickness of approximately 20 μm of electroformed palladium, while the centre-to-centre spacing between individual needles was 200 μm. Besides, the use of palladium as a structural material provided high mechanical strength and durability, as well as, biocompatibility for use in biomedical applications. By means of the same process, Chandrasekaran et al. [118] proposed microneedles fabricated by using electroplated metals including palladium, palladium-cobalt alloys and nickel. Particularly, single microneedle and multiple microneedle arrays showed features enabled by the process, such as complex tip geometries, micro barbs, mechanical penetration stops and multiple fluid output ports. The microneedles length was 500 µm, 1000 µm and 1500 µm. In addition to the needle shaft, each needle contained a sharp tip ranging in length from 250 µm to 700 µm depending on the tip taper, while the cross-section ranged from 70 μm to 200 μm in width and 75 to 120 μm in height, with a wall thickness of about 30 μm. The microneedle arrays were typically 9.0 mm in width and 3.0 mm in height with between 3 and 17 needles per array.
A different method for realizing hollow in-plane microneedles was proposed by Parker et al. [124]. The main difference respect to previous process is that a sacrificial layer is not necessary, and the inner channel was obtained by bonding a microneedle array with an unpatterned foil. The starting material was a titanium foil, where the planar geometry of the needles and their fluidic channels was defined by means of patterning a mask layer. Afterwards, selective etching steps allowed microneedle fabrication and channels realization on microneedle surface. In order to close the microneedle channels and obtain a hollow microneedle arrays, a gold thermo compression bonding to an unpatterned titanium foil was performed. Finally, by means of a selective etching of the titanium foil using the patterned substrate as an etch mask, titanium hollow in-plane needles resulted. The linear arrays presented were composed of 10 microneedles of three different lengths (500 μm, 750 μm, and 1000 μm) fabricated with two 25 μm thick titanium foils. All microneedles were 100 μm wide with an internal channel 30 μm wide and 10 μm high.
1.5.3.2 Metallic array of out-of-plane microneedles
Out-of-plane hollow microneedles [121, 122, 125, 126, 128] are generally fabricated by means of a conformal electroplating of metals on a micromachined mold containing the desired needles structure (fig. (1.7.c)). Particularly, the mold could be a replica of the needles or a complementary structure (as shown in figure (fig. (1.7.c))). The mold could made out of a metal (i.e. titanium) [121] a polymer
