IEICE TRANS. ELECTRON., VOL.E95–C, NO.2 FEBRUARY 2012188

INVITED PAPER

Special Section on Photonic Devices using Nanofabrication Technology and Their Applications

Photonic Crystal Nanolaser Biosensors
Shota KITA†,††a) , Shota OTSUKA†,†† , Shoji HACHUDA†,†† , Tatsuro ENDO††† , Yasunori IMAI††† ,Yoshiaki NISHIJIMA†††† , Hiroaki MISAWA†††† , Nonmembers, and Toshihiko BABA†,†† , Member

SUMMARY
High-performance and low-cost sensors are critical devices for high-throughput analyses of bio-samples in medical diagnoses
and life sciences. In this paper, we demonstrate photonic crystal nanolaser
sensor, which detects the adsorption of biomolecules from the lasing wavelength shift. It is a promising device, which balances a high sensitivity,
high resolution, small size, easy integration, simple setup and low cost. In
particular with a nanoslot structure, it achieves a super-sensitivity in protein sensing whose detection limit is three orders of magnitude lower thanthat of standard surface-plasmon-resonance sensors. Our investigations indicate that the nanoslot acts as a protein condenser powered by the opticalgradient force, which arises from the strong localization of laser mode in
the nanoslot.
key words: photonic crystal, nanolaser, nanoslot, liquid sensor, biosensor

1.

Introduction

Protein analyses are increasing demands for medical diagnosis, health care, food evaluation, proteomics, and so on.
In particular, high-throughput and low-cost assay of a target
protein in a small amount of analyte is desired and studied [1]. Lab-on-a-chip and micro-total analysis systems (μTAS) [2] have been developed as supplementary tools that
control fluids and promote the reactions in the target. Similarly, developing novel principles and methods in sensors is
becoming an important issue [3]. At present, labeling methods are commonly used for sensing. Here, fluorescent molecules are functionalized as labels to the target so that the
behaviors of the target is observed directly through versatile
instruments such as fluorescent optical microscopes. However, the complicated functionalization process is a bottleneck for high throughput and cost reduction. In addition, thefunctionalization itself remains an uncertainty if it maintains
the original behaviors of the target.
As alternatives, various label-free methods have beenstudied and developed. Fundamentally, they are required tosatisfy the following requirements:
1) High performance: The most fundamental spec of bioManuscript received July 1, 2011.
Manuscript revised November 1, 2011.
†
The authors are with Yokohama National University,
Yokohama-shi, 240-8501 Japan.
††
The authors are with CREST, Japan Science and Technology
Agency, Tokyo, 105-0075 Japan.
†††
The authors are with Tokyo Institute of Technology,
Yokohama-shi, 226-8502 Japan.
††††
The authors are with Hokkaido University, Sapporo-shi, 0010021 Japan.
a) E-mail: d09sd103@ynu.ac.jp
DOI: 10.1587/transele.E95.C.188

sensors is the detection limit (DL) meaning the minimum
detectable concentration of the target in the analyte. It
is determined by the sensitivity of the sensor and also
by the resolution of the sensor and other equipments,
which is aﬀected by noise. For high-throughput assay
simply checking the presence of the target without preconcentration process, a low DL of the order of or evenlower than 1 pM is necessary [1] (DL of labeling methods is usually of nM order and sometimes reaches to pMorder in researches). The quantization of the target issometimes required in medical applications, and a wide
dynamic range showing a linear response is necessary
for this. Of course, the stability of results and insensitivity to the measurement condition reduce the noise andimprove the reliability and usability.
2) Low cost: To avoid the mutual contamination of samples, the disposable use of sensors is preferable, andtherefore, the sensor cost is crucial. In addition, hours
of work for sensing also reflect to the total analysis cost.
Regarding this point, neglecting the functionalization oflabels is already advantageous. A low DL allowing skipping the pre-concentration process adds another advantage.
3) Functionality: If we can monitor the realtime responseof the sensing, it is useful for analyzing the kinetics ofreaction. If the multi-information can be obtained simultaneously for multi-targets, it enhances the throughput
and enables statistical analyses.
4) Usability: A simple and portable system is desirable sothat anyone can use anywhere. For this purpose, downsizing the total system is necessary, and this also contributes to the integrability. The disposable use is also
important for this.
5) Integrability: If sensors are integrated with μ-TAS, it will
be a powerful inspection kit that achieves high performance in terms of throughput, stability and functionality.
For this purpose, sensors must have a size comparable toor even smaller than fluidic channels in μ-TAS, whichusually have a feature size of 10 μm order.
Table 1 summarizes the performance and other features of label-free biosensors reported so far. Fundamentally, all of them use the same procedure of the sensing.
First, a host material known to bind specifically with the target guest protein is functionalized onto the sensor surface.
Then, the sample containing the target is injected and thec 2012 The Institute of Electronics, Information and Communication Engineers
Copyright KITA et al.: PHOTONIC CRYSTAL NANOLASER BIOSENSORS

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Table 1Comparison of label-free biosensors.

specific binding is detected by the sensor. Such two step
process is much simpler than labeling methods including
additional functionalization of labels. The most common
sensor is the surface plasmon resonance (SPR) sensor [4].
It measures the resonant angle of incident light coupled tothe surface plasmon mode of a metal, and detects its shift
against the presence of the target on the metal. The system
is easy to use but hard to miniaturize, and its DL is on anaverage level due to the broad resonant spectrum, which degrades the resolution. The local plasmon resonance at metal
nanoparticles [5] and surface-enhanced Raman scattering at
rough metal surfaces [6] achieve an ultralow DL through thehighly sensitive spectral shift, but the signal is noisy and unstable. Field eﬀect transistors (FETs) modified by carbon
nanotubes (CNT) [7], [8] or graphene [9] can detect the specific binding through the change of the electric characteristics. Such one chip device is easy to handle, but may be too
expensive for disposable use. In addition, they do not show
a good DL and robust signal simultaneously.
Optical micro/nanocavities are also exploited as sensors [10]. Here, tunable laser light is coupled to the device,
the resonant spectrum is measured, and the binding is detected through the spectral shift. Cavities are categorized
into the whispering gallery mode type (sphere [11], toroid[12], ring [13], [14]) and the Bragg reflection type (grating[15], photonic crystal (PC) [16], [25]). Microspheres andmicrotoroids recorded an excellent DL thanks to ultranarrow spectral linewidths of less than 1 pm. However, sucha narrow spectrum requires complicated measurement tools
and leads to unwanted response against environmental fluctuations. Low cost manufacturing of microring and PC cavities will be possible by exploiting the CMOS-compatible
technology in Si photonics [17], but their DL is not so good
because of the relatively broad spectrum and moderate sensitivity. In addition, all of these passive cavities need theoptical fiber coupling at input and output (I/O), which becomes a barrier for cost reduction and disposable use.
In this paper, we focus on an alternative method using lasers (in other words, active cavities). The fundamental
principle is similar to that of passive cavities; the binding
is detected through the lasing spectral shift. Standard stripe
lasers and vertical cavity surface emitting lasers may also
be applicable [18], but a high sensitivity is not expected as
their laser mode is almost confined inside the semiconductors. On the other hand, micro/nanolasers have laser modes
whose large fraction of fields is penetrating to air, resulting
in a high sensitivity [19]. In addition, they have the following advantages in comparison with passive cavities:

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1901) A high resolution is ensured by a narrow spectrum above
the laser threshold. This eliminates a general constraint
in passive cavities that a smaller cavity degrades its Q
and expands the spectrum. Thus, nanolasers can be
downsized independently of the passive Q below thethreshold if Q is suﬃcient for lasing. The lasing also
eliminates the instability of passive Q easily aﬀected by
the disordering.
2) Optical I/O are much more simplified than the case ofpassive cavities, since both the photopumping and thedetection of laser emission can be done through freespace optics. Therefore, the sensor chip can be isolated
from other measurement equipments. The chip cost willbe almost negligible even though it is fabricated by using
e-beam lithography, because the lithographic area is only(∼10 μm)2 for each chip. This allows the manufacturingof low cost disposable chips using a standard semiconductor process. Assuming two state-of-the-art e-beam
writers, the throughput can be >14,000 chip/day. The
chip cost can be much less than Y=100, even including
costs for other facilities, materials, packaging, testing,staﬀs and yield.
3) The spectral shift for nanolasers is measured by a spectrometer such as optical spectrum analyzer (OSA). Here,a high-speed spectral analysis is possible by using sensor arrays and mixing coarse and fine observations withchanging the span and resolution. Such high-speed flexible measurement is not easy for passive cavities when ahigh Q mode is searched by scanning the tunable laser
source; the rough scanning can skip the cavity mode.
The performance of nanolaser in this work, i.e.
nanoslot (NS) photonic crystal (PC) H0-type nanolaser[20], [21], is compared with other biosensors in Table 1.
Since target proteins are diﬀerent between reports, the aﬃnity constant KA changes from 105 − 1011 M−1 . In general,
DL is improved by a large KA . Therefore, the performance
should be evaluated by the product of DL and KA . Let us
define the figure-of-merit (FOM) factor as (DL × KA )−1 normalized by that for the SPR sensor reported recently [22].
Then our nanolaser achieves a very high FOM of 230 −
2,300 assuming the intensity noise factor ξ, which will be
discussed in Sect. 3, to be 1.0 − 0.1, respectively.
To date, FOMs higher than these value have only
been reported for Au nanoparticles [6] and micro-toroidal
cavity [12], which have aforementioned drawbacks. Thisnanolaser is very unique because a large fraction of lasermode is strongly localized in the NS outside of semiconductors. This paper reports in detail later that this localization improves the thermal characteristics and resolution, andgreatly enhances the sensitivity for biomolecules, resulting
in the very low DL and high FOM. In Sect. 2, we first describe the structure and laser characteristics of this device.
We particularly focus on eﬀects of the NS for the thermal
characteristics and resolution. In Sect. 3, we present the theoretical sensing characteristics and experimentally demonstrate the sensing of liquids. In Sect. 4, we show biosensing

experiments for nonspecific and specific bindings, including realtime responses. We discuss the mechanism of thesuper-sensitivity arising from the NS.
2.

NS PC Nanolaser

The PC H0 nanolaser without NS consists of the shiftof nearest two or four airholes in a triangular lattice PC
slab. The three-dimensional finite-diﬀerence time-domain(FDTD) simulation predicts that a device fabricated into
GaInAsP quantum-well wafer with a lattice constant a =500 nm, airhole diameter 2r = 260 nm, s x = 120 nm,
sy = 80 nm, and slab thickness t = 200 nm has a passive Q ∼ 100,000 and a modal volume Vm ∼ 0.15(λ/n)3 ,where λ is the wavelength and n is the material index at
the mode maximum. We have already reported its roomtemperature continuous-wave (cw) operation with a subμW threshold [21], [26], high-resolution liquid index sensing (∼9.0 × 10−5 RIU) [21], [27] and protein sensing [28].
However, a broad spectral linewidth caused by the large
thermal chirping [29] have degraded the sensing resolution[30]. To solve this problem, we incorporated the NS, which
maintains the localized mode, as shown in Fig. 1. Similarmode has first been discussed in Ref. [31], and it has been
exploited for reducing Vm in a passive cavity [32], capturingbiomolecules such as DNAs [33] and enhancing the liquidindex sensitivity [34]. For a high-index-contrast interfaceof the NS, Maxwell’s equations dictate that, in order to sat-

Fig. 1 NS nanolaser. (a) Scanning electron micrograph (SEM) of fabricated device. Details of fabrication are shown in Ref. [21]. (b) Calculated
modal distribution for a nanoslot width wNS = 20 nm. Left and right panels
depict in-plane and cross-sectional distribution, respectively. Dashed line
shows the location of cross-section.

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Fig. 2 Laser characteristics of NS nanolasers in water. Pirr shows irradiation power of photopumping. Insets show near field pattern of lasing
spectra and near field pattern.

isfy the continuity of the normal component of the electric
flux density D, the corresponding E field must undergo alarge discontinuity with a much higher amplitude in the lowindex side. This discontinuity strongly enhances the field
and forms a localized mode in the low-index NS. In the caseof NS nanolaser, it increases the total mode penetration outside of semiconductors by 10–20%. Since the biosensing is
usually performed in water-based analyte, the mode penetration enhances the sensitivity. Note that water has an absorption coeﬃcient of ∼10 cm−1 at λ ∼ 1.55 μm and a local
maximum of ∼30 cm−1 at l.45 μm [35]. It is a constraint insetting the modal wavelength, while a suﬃciently high Q of
10,000 is theoretically expected at λ ∼ 1.55 μm even withthe absorption, maintaining an extremely small Vm = 0.006(λ/nwater )3 when s x and sy in Fig. 1(a) are optimized [21].
With sensor applications in mind, the laser characteristics were experimentally evaluated in water using pulsed
photopumping at 980 nm (duty-ratio of 200, spot diameterof 20 μm for eﬃcient pumping of the NS mode [21]) andthe laser emission was detected using a fluorescence microscope setup. The laser spectrum was analyzed by OSA witha resolution of 10 pm. Figure 2 shows the results for several
devices with a NS width wNS of 30–60 nm. Here, the eﬀective pump power Peﬀ was estimated by counting the spatial
overlap of the pump spot with the eﬀective modal area andits imperfect absorption in the slab [21]. Clear lasing is observed with a 35 dB peak intensity at an eﬀective thresholdof 3–5 μW. Figure 3 compares the lasing spectra with andwithout the NS. As mentioned above, the spectrum without
NS under pulsed condition widely broadens due to the thermal chirping. The full-width at half-maximum (FWHM) of
the spectrum, Δλw , is over 600 pm in water, which is 15
times narrower than in air [29] but still broad. In contrast,
it narrows drastically to 18 pm with the NS, which is close
to the resolution limit of the OSA. Since water has a negative thermo-optic coeﬃcient Δn/ΔT ∼ −10−4 RIU/K [36]
(RIU means the refractive index units), the positive coeﬃcient ∼10−4 RIU/K of GaInAsP slab can be canceled and theFig. 3 Comparison of pulsed lasing spectra in water between with andwithout NS.

Fig. 4 Comparison of wavelength shift due to (a) temperature and (b)irradiated power change between with and without NS.

thermal chirping can be suppressed by the appropriate mode
penetration into the NS. This expectation was confirmed further by measuring the temperature dependence of the spectral shift Δλ s with and without the NS, as shown in Fig. 4(a).
Without NS, the shift increases almost linearly with temperature with a slope of 82 pm/K. With the NS, on the other
hand, the dependence is reduced to four times smaller at T
= 20◦ C and almost eliminated (< 3 pm/K) at 40◦ C. This temperature dependence of athermalization can be explained by
the quadratic thermo-optic coeﬃcient of water [36]. Solidlines in Fig. 4 are obtained by taking account of these effects with the modal profile in Fig. 1. They almost explain
the experimental behaviors. This athermalization also reflects to the power dependence of the spectral shift. It is
reduced to six times smaller with the NS at T = 20◦ C, as
shown in Fig. 4(b).
3.

Sensing Characteristics

The nanolaser sensor detects the index change within thepenetration depth of the laser mode (1/e depth is calculated
to be 110–140 nm for the above devices) through the spec-

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tral shift. The index resolution Δnres is given by

−1
Δλ s
Δnres = ΔλminΔn(1)

where Δλmin is the minimum resoluble shift and Δλ s /Δn is
the sensitivity. Let us assume a Lorentzian spectrum I(λ)
with the peak intensity I peak and peak wavelength λ p , thewavelength shift amount Δλ and the fluctuation noise in the
wavelength δλ limited by the stability of device and the accuracy of OSA, and intensity δI caused by the temporalchange of pumping power and degradation of the device.
The intensity at the intersection point between λ and λ + Δλis denoted as I(λ p + Δλ/2). The wavelength shift Δλ is detectable, when I peak − I(λ p + Δλ/2) is higher than δI (Δλmin
is determined when they are equal). Solving this condition,
adding δλ and assuming I peak > δI, Δλmin is approximated
as

1
(2)
Δλmin = ξΔλw + δλ, ξ =I peak /δI − 1
ξ is the intensity noise factor showing how the intensity
noise aﬀects the wavelength fluctuation based on the assumed I(λ). In nanolasers under sensing, δI could be large
because of the convection of water leading to the fluctuation in photopumping and light extraction condition. WhenI peak /δI = 2, 5 and 10, the intensity noise factor ξ becomes
1.0, 0.5 and 0.3, respectively. If data acquisition is suﬃciently averaged by the statistic measurement, I peak /δI = 100
and ξ  0.1 are obtainable. In the OSA used in this study,
the wavelength reproducibility is < 6 pm. The temperaturefluctuation can be 0.5 K for long measurements, and δλ becomes < 15 pm and ∼400 pm with and without NS, respectively. The noise induced from the pump source is smaller
than this. Then, the sum of all the noise approximately gives
Δλmin ∼ Δλw [21]. In realtime continuous measurements,
short term fluctuations in temperature and OSA are suﬃciently small. Thus δλ  Δλw , and Δλmin ∼ ξΔλw where ξ 0.1 is possible.
On the other hand, the sensitivity Δλ s /Δn in the NS
nanolaser is calculated by using the FDTD method, as
shown in Fig. 5(a). Here, the uniform bulk medium (theindex nb ) simulates a liquid and the 20-nm-thick film(n f ) on surfaces and sidewalls simulates the adsorption of
biomolecules. The modal wavelength in water is used as areference. The sensitivity for the film Δλ s /Δn f is 108 and82 nm/RIU with and without the NS, respectively. Since themodal field overlaps with the film partially, the sensitivity isnearly four times smaller than that for bulk, Δλ s /Δnb . The
field component localized in the NS enhances Δλ s /Δn f by
25%. Figure 5(b) shows a similar calculation assuming thefilm only inside the NS. The spectral shift is smaller than (a),but a shift around 1 nm in this figure is still large enough todetect and resolve in our setup. The sensitivity is enhanced
in inverse proportion to wNS . While biosensing experiments
are shown in the next section, Fig. 6 demonstrates the liquid index sensing. The sensitivity is obviously enhanced

Fig. 5 Spectral shift calculated with index change, where modal wavelength in water with nb = 1.321 is used as a reference. (a) Comparison
between bulk and film (20 nm thickness) index sensitivity. (b) Result forfilm only in NS.

Fig. 6 Spectral shift in diﬀerent index liquids Cargille B0700/0701.
Spectral widths are diﬀerent from those in Fig. 3 because of diﬀerent pumping condition and diﬀerent thermo-optic coeﬃcients of index liquids.

from 280 nm/RIU without NS to 410 nm/RIU with the NS.
These values are in good agreement with the calculation in
Fig. 5(a).
4.

Biosensing

As mentioned earlier, label-free biosensing uses the twostep process, i.e. the host functionalization and guest adsorption. The fundamental procedure of this experiment
including this process is depicted in Fig. 7. First, a hydroxyl group is functionalized on the device surface with 9%
HCl/deionized water at 3◦ C for more than 2 hours. Then,KITA et al.: PHOTONIC CRYSTAL NANOLASER BIOSENSORS

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Fig. 7

Schematic procedure of chemical treatment and biosensing.

Fig. 8 Sensing of BSA protein in the high concentration regime. (a) Wavelength shift before and after
BSA adsorption. (b) Dependence of wavelength shift on NS width. Solid lines denote averages.

silanization is performed using 10% N-2-(aminoethyl)-3aminopropyltrimethoxysilane (KBM-603, Shinetsu Chemical Co.) and a host material is subsequently functionalized
at 20◦ C for 1 hour. Finally, a target guest is adsorbed andthe nanolaser detects it. After soaking the device in highconcentration (several μM) guest aqueous solution for morethan 30 minutes, we usually observe approximately 10-nm
roughness on the device surface and in the NS using SEM,
which is not observed before the process. This indicates theadsorption of the guest protein. In this experiment, we chose
two combinations of host material and target guest material.
One combination is glutaraldehyde (GA) and bovine serum
albumin (BSA, molecular weight = 68 kDa) [37]. BSA is acommon protein for test experiments in biosensing. It binds
with GA nonspecifically with a relatively small aﬃnity constant (∼107 M−1 order). The other combination is biotin andstreptavidin (SA, molecular weight = 53 kDa) [38]. This
combination is often used as a model of host-guest specificbinding. The aﬃnity constant in this case is very large (>1013 M−1 ) [39].
Figure 8(a) shows laser spectra before and after the
BSA adsorption when the BSA concentration was as highas 10 μM. The spectrum clearly redshifted for both with andwithout the NS. Without the NS, the spectrum is broadenedby the thermal chirping but the shift is still observed clearly.
With the NS, the spectrum is markedly narrowed and theresolution is drastically improved. Figure 8(b) summarizes
the dependence of the redshift on the NS width. The average

shift with and without the NS are 4.4 nm and 2.2 nm, respectively. It means that a twice higher sensitivity was obtainedby the NS. This enhancement is larger than that expectedin the theoretical result of Fig. 5(a). This suggests that the
NS not only improves the sensitivity but also accelerates theadsorption of BSA, which is further discussed below.
Figure 9(a) shows the spectral shifts for diﬀerent BSA
concentrations C. A large shift ShiftH is observed in thehigh-concentration regime of >10 nM for both with andwithout the NS. It is not saturated even at > 100 μM. Thismay be due to the physisorption, aggregation and phase transition [40] of high-concentration BSA. On the other hand, asmall shift ShiftL is observed clearly in the low concentration regime only with the NS. This was observed repeatedly
for all devices in diﬀerent fabrication lots. Therefore, we
can consider that ShiftL arises from the adsorption inside theNS, while ShiftH from the adsorption over all other surfaces.
Actually, we observed after the experiment that BSA is selectively adsorbed in/around the NS even at C = 1 pM, as
shown in the inset of Fig. 9(b). Usually, such dense adsorption does not occur in a low concentration regime. Provided
that each adsorption process occurs independently, Δλ s isexpressed as the sum of three diﬀerent Langmuir adsorption
isotherms [41]:Δλ s =

Δλmax1 KA1C Δλmax2 KA2C Δλmax3 KA3C++
1+KA1C
1+KA2C
1+KA3C(3)

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194

Fig. 9 Wavelength shift with BSA concentration. (a) Spectra for diﬀerent concentrations. (b) Peakwavelength shift with concentration. Circular plots show experimental data averaged over many devices
with wNS = 30–60 nm and error bars show data scattering as well as spectral fluctuation. Fitting curves
are obtained from Eq. (3). The inset shows magnified SEM picture around the NS afters BSA adsorption
at 1 pM.
Table 2Fitting parameters for solid lines in Fig. 9(b) and evaluated DL assuming ξ = 1.0.

chemisorption inside and outside of the NS, respectively,
and KA3 is that of the physisorption. Only KA1 is influenced
by the presence of the NS. Δλmaxi (i = 1, 2, 3) is the maximum shift when each adsorption is saturated. Figure 9(b)summarizes Δλ s of many devices with C. Theoretical linesof Eq. (3) fit well with experimental plots when parametersin Table 2 are assumed. Note that KA1 determined from thefitting is much larger than KA2 . This means that BSA istrapped around the NS and its adsorption is particularly accelerated. The ratio of KA1 to KA2 gives the trapping potential Utrap [42] such that
Utrap = kB T ln(KA1 /KA2 )(4)

where kB T is the thermal energy. From the parameters in
Table 2, Utrap is calculated to be 9.7kB T . We also estimated that the mode localization evaluated from pump anddetected powers and various losses in the optical setup gives
rise to the optical gradient force, whose trapping potential is
4–20kB T . Thus, this force is a possible trapping mechanism.
Other mechanisms from pump light and thermal eﬀects arealso worth considering, because they could explain the continuous supply of dispersed BSA to the NS. The strong field
localization near the two joint areas between airholes andNS, as shown in Fig. 1(b), may cause a specific thermal gra-dient, which can also be a mechanism of the protein trapping. In any case, DL is greatly improved by the enhanced
KA1 . The DL is equivalent to C at Δλ s = Δλmin , and so derived as−1
DL = KAi
Δλw /(Δλmaxi − Δλw )(5)

By substituting parameters in Table 2 into Eq. (5), DL without NS is estimated t o be 38 nM (= 2.6 μg/ml for BSA),
while that with the NS, 255 fM (= 17 pg/ml). The realtime response of the spectral shift reflects the dynamic diffusion and adsorption of BSA. To observe this, we intermittently injected the solutions of diﬀerent BSA concentrations
from low to high using simple fluidic channels composedof glasses, plastics or PDMS polymer. All samples weremanually injected by using a micro-pipette. Figure 10(a)focuses the response (sensorgram) in the low-concentration
regime. Discontinuities just after injections were caused by
the defocusing of pump light, which temporally stopped thelasing. We can observe a clear shift of 0.1–0.2 nm between
5 and 24 pM, which corresponds to ShiftL in Fig. 9(b). Although it is small at the injection of 36 pM, the response
certainly exhibits a shift. From 36 pM to 1.6 nM, the irregular step shift and gradual shift are mixed. It is considered
to be due to the nonuniform distribution of BSA in the low-

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195

Fig. 10 Sensorgrams for (a)–(d) adsorption and (e), (f) desorption of BSA. (a) and (b) exhibit the
step-like shift due to intermittent injection of diﬀerent concentration BSA for low and high concentration regimes, respectively. (c)–(f) show sensorgrams fitted with theoretical curves calculated with
the Langmuir model considering diﬀusion of protein [43]. Used fitting parameters are denoted in eachfigure. (c) compares two fitting curves.

concentration regime and/or partially saturated adsorption at
the NS. In the high-concentration regime, on the other hand,the step increase is observed more clearly in Fig. 10(b). The
response is slower than the case in the low-concentration
regime. To analyze the adsorption and desorption kinetics of BSA, we fitted theoretical curves, which were calculated with the Langmuir model considering the diﬀusionof protein [43], with the observed sensorgrams under different situations. In the calculation, we also assumed fast
protein diﬀusion such as 100 μm/s, considering the force ofthe direct sample injection. Figure 10(c) shows the sensorgram in the low concentration regime. Here, two theoretical curves were calculated for diﬀerent dissociation rate
constants kd of 1.0 × 10−4 s−1 and 3.0 × 10−3 s−1 , which

simulate the desorption from thin- and dense-adsorbed BSA[44], respectively (the adsorption rate constant ka is given by
KA kd ). Then, we optimized KA so that the theoretical curvefits to the experimental plots. It is seen in Fig. 10(c) thatthe round curve for larger kd fits better. The corresponding
KA is 3.5 × 1011 M−1 , which is almost comparable to KA1 in
Table 2 estimated for NS devices. Thus all the results consistently show that the dense adsorption actually occurred
around the NS even in the low concentration regime. Figures 10(d) and (e) show the adsorption and desorption in
the high concentration regime, which should reflect the normal chemisorption between GA and BSA. Both theoreticalcurves are in good agreement with experimental plots bysetting the same ka and kd . Here, KA2 = 3.3 × 107 M−1 was

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Fig. 11 Sensorgrams of biotin-SA specific binding (red circle plots), which is compared with physisorption of SA (black circle plots). Red, black and green curves depict theoretical fits to biotin-SA
specific binding, SA physisorption and the sum of both adsorption processes, respectively.

estiamted from the curve fitting, which is also on the sameorder as that in Table 2. The desorption behavior for thephysisorption is also confirmed from Fig. 10(f), and a KA3
similar to that in Table 2 was also estimated.
We also tested the realtime sensing for the specificbinding of biotin-SA, as shown in Fig. 11. Since such
realtime specific binding of biomolecules has never been
observed with nanolasers, we used high-concentration SA(2.8 μM) in this preliminary experiment, and the result wascompared with the case without biotin. After injecting SA,
both cases exhibit the redshifts. The small redshift without
biotin might be caused by the physisorption of SA on thedevice surface. The biotin-SA binding shows eight times
larger shift, indicating that the specific binding is more
dominant than the physisorption. Similarly to the resultsin Fig. 10, we can obtain fitting curves assuming KA andkd . When we only assume KA = 1.9 × 1013 M−1 for thebiotin-SA specific binding in Ref. [39] and a very small kd
of 1.0 × 10−7 min−1 , the red curve in Fig. 11 was obtained,
which does not fit to the experimental plots. It might be influenced by the physisorption of SA simultaneously occurring due to the rapid increase in the SA concentration. Indeed, the green curve calculated from the sum of these twoadsorption processes fits much better. The slight disagreement could be caused by the small diﬀerence of diﬀusion
speed for each manual injection.
5.

Conclusion

Nanolasers well satisfy the requirements of biosensing interms of performance, cost and usability. By using photopumped nanolaser as a sensor chip isolated from other equipments, its cost reduction and disposable use are possible.
We fabricated GaInAsP PC H0 nanolasers with a nanoslotof 30–60 nm width as a sensor chip. We observed the laser
operation with a high yield even in water, which is suitablefor biosensing. In particular, the nanoslot athermalized thenanolaser in water and achieved a narrow spectrum, whichis eﬀective for high-resolution sensing. We applied this device to the BSA protein sensing and unexpectedly evaluateda super-sensitivity. It gives an ultralow detection limit of
255 fM even assuming the worst intensity noise. This value
is three orders of magnitude lower than that of SPR sensors
reported recently [22]. The super-sensitivity was explainedby the trapping of BSA in the nanoslot, to which the opticalgradient force of localized laser mode contributes. We also
demonstrated the detection of the biotin-streptavidin binding as a preliminary experiment toward future specific detection of biomarkers in medical diagnoses. Further investigations of the trapping mechanism may lead to a unique control tool for manipulating and detecting biomolecules more
eﬃciently.
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IEICE TRANS. ELECTRON., VOL.E95–C, NO.2 FEBRUARY 2012198

Shota Kita
received the B.E. and M.E.
degrees from the Department of Electricaland Computer Engineering, Yokohama National
University, Yokohama, Japan, in 2007 and 2009,
respectively. He is currently working towards
the Ph.D. degree in photonic crystal nanolasers
and their applications for liquid and biosensing
at the same university. Mr. Kita is a memberof the Japan Society of Applied Physics. He
received two academic awards including JSAP
presentation award in 2010.

Shota Otsuka
received the B.E. degreefrom the Department of Electrical and Computer Engineering, Yokohama National University, Yokohama, Japan, in 2011. He is currently
working towards the M.E. degree in ultra highsensitivity photonic crystal nanolaser biosensor
at the same university. Mr. Otsuka is a memberof the Japan Society of Applied Physics.

Shoji Hachuda
received the B.E. degreefrom the Department of Electrical and Computer Engineering, Yokohama National University, Yokohama, Japan, in 2010. He iscurrently working towards the M.E. degree
in bio-electronics and surface chemistry andphotonic crystal nanolaser biosensor for biomarker protein detection at the same university.
Mr. Hachuda is a member of the Japan Societyof Applied Physics.

Tatsuro Endoreceived his Ph.D. degreefrom the Japan Advanced Institute of Scienceand Technology (JAIST), Biotechnology Laboratory in the Dept. of Biological Science andBiotechnology, in 2006. He was an assistantprofessor at P&I Lab of Tokyo Institute of Technology in 2006–2011. Then he became an associate professor at the Department of Applied
Chemistry, Osaka Prefecture University. His
current field of interest includes nanophotonics,
biophotonics and nanobiotechnology.

Yasunori Imai
received the B.E. degreefrom the Department of Bioengineering, Tokyo
Institute of Technology, Japan, in 2010. He iscurrently working towards the M.E. degree in
SiN photonic crystal and their applications forsensing bio-molecules at the same university.

Yoshiaki Nishijima
received the B.E.
degree from the Facility of Science Divisionof Chemistry, Hokkaido University, in 2004,
M.D. degree from Graduate School of Science,
Hokkaido University in 2006, and Ph.D. degreefrom Graduated School of Information Science,
Hokkaido University in 2009. During his Ph.D.
work, he had been engaged in fabrication andcharacterization of opal photonic crystals. From2009 to 2010, He worked as a postdoctrate fellow at Research Institute for Electronic Science,
Hokkaido University, working on the plasmonic solar cell. He became an
assistant professor at Yokohama National University in 2011. His recent
study is focusing on material science and chemistry especially for visible
to THz wavelength plasmonic materials. Dr. Nishijima is a member of theJapan Society of Applied Physics and Chemical Society of Japan.

Hiroaki Misawa
is Director and Professor
at the Research Institute for Electronic Science,
Hokkaido University, Sapporo, Japan. He obtained his B.Sc. in Engineering from the Department of Engineering at Tokyo Metropolitan University in 1979, graduated in Engineering fromthe University of Tsukuba in 1981 and completed his Ph.D. at the Department of Chemistry
there in 1984. After postdoctoral positions at the
Universities of Texas at Dallas and an AssistantProfessorship at the University of Tsukuba, hejoined the Department of Engineering of University of Tokushima in 1993
and was promoted to full professor in 1995. His current research interests include light-matter interaction, ultra-fast processes in materials, plasmonics, and strong photon-molecule coupling fields for chemical reaction.
He authored more than 260 papers. He was the representative researcher
of Core Research for Evolutional Science and Technology by Japan Science and Technology Corporation from 2003 to 2007. He had engaged in
Grant-in-Aid for Scientific Research (KAKENHI) on Priority Area entitled “Strong Photon-Molecule Coupling Fields for Chemical Reactions”,
MEXT, Japan from 2007 to 2011. He is Deputy Editor of the Journal of
Photochemistry and Photobiology C, Photochemistry Review and a member of ACS Journal of Physical Chemistry C Advisory Board.

Toshihiko Baba
received the Ph.D. degree from the Division of Electrical & Computer Engineering, Yokohama National University (YNU), Japan, in 1990. During his Ph.D.
work, he had been engaged in ARROW waveguides and photonic integration. From 1991–
1993, he joined Tokyo Institute of Technology as a research associate and investigated thespontaneous emission control and room temperature cw operation in VCSELs. He became an
associate professor and full professor of YNU
in 1994 and 2005, respectively. In these 18 years, he has studied photonic nanostructure devices including photonic crystals (PCs), microdisks,
and Si photonics devices. His recent topics are regarding slow light PC
waveguides towards tunable delay, nonlinear enhancement, and high-speed
modulators, Si photonics devices and circuits based on CMOS-compatible
process, and bio-sensing using PC nanolasers. He is an author of 150 journal papers with more than 5200 citations. Dr. Baba is a member of IEICE,
JSAP, IEEE/Photonics, OSA, and APS, and an associate member of Science Council of Japan. He received 10 academic awards including JSPS
Award in 2005, IEEE/LEOS Distinguished Lecturer Award in 2007, and
IEICE Electronics Society Award in 2011.