JOURNAL OF LIGHTWAVE TECHNOLOGY, VOL. 39, NO. 12, JUNE 15, 2021
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Cancer Biomarker Detection With Photonic
Crystals-Based Biosensors: An Overview
Alberto Sinibaldi
(Invited Paper)
Abstract—This review wants to give an overview of the photonic crystals-based biosensors for cancer biomarkers detection.
Indeed, in the last two decades, 1D, 2D, and 3D photonics crystals have seen an extraordinary development in the direction of
medical diagnosis, health assessment, and therapy monitoring.
Cancer-related biomarkers can span over a wide range of biological elements including circulating tumor DNA, miRNA, proteins,
enzyme, metabolites, as well as circulating tumor cells. Therefore,
the review is articulated in three sections reporting on the basics
of the most common used 1D, 2D, and 3D photonic crystal configurations followed by the more recent biosensing applications in
cancer biomarker detection. These devices include 1D truncated
multilayers such as distributed Bragg reflectors and layered gratings, 2D ordered waveguiding slabs with particular emphasis on
the micro/nano cavities, and 3D direct and inverse opals. Their
added value can be resumed in the capability to strongly confine
the electromagnetic radiation interacting more efficiently with the
biological sample thus improving the limit of detection. In conclusions, photonic crystal-based biosensors hold great potential
in the detection of cancer biomarkers thanks to their ultimate
performances guarantying, in the near future, a versatile sensing
tool to clinical personnel and physicians.
Index Terms—Cancer biomarkers, fluorescence biosensing,
label-free biosensing, optical biosensors, photonic crystals.
I. INTRODUCTION
HOTONIC Crystals (PCs) based biosensors continue to be
high on the agenda because of the ever-increasing need to
face new challenges in healthcare and precision medicine. Indeed, in recent years, such types of biosensors have been largely
exploited for the specific and precise quantification of various
biomarkers in complex biological media enhancing sensing performance. Thanks to the periodic structuring of the constituent
materials, in the majority of the cases obtained artificially, PC
can allow or stop the propagation of the light with wavelength in
the order of their periodicity. Thereby, a photonic band structure
is generated in the optical transmission spectrum by prohibiting
the transport of the photons in certain energy ranges called
P
Manuscript received December 15, 2020; revised January 28, 2021; accepted
January 29, 2021. Date of publication February 2, 2021; date of current version
June 16, 2021. This work was supported in part by Italian MIUR Ministry (NeoN,
IDARS01_00769) and in part by Regione Lazio (TURNOFF, 85-2017-14945).
The author is with the Department of Basic and Applied Science
for Engineering, SAPIENZA University of Rome, 00161 Rome, Italy
(e-mail: alberto.sinibaldi@uniroma1.it).
Color versions of one or more figures in this article are available at https:
//doi.org/10.1109/JLT.2021.3056225.
Digital Object Identifier 10.1109/JLT.2021.3056225
photonic bands gap (PBG) [1], [4]. Thus, in PC, the dielectric
function or the refractive index (RI) is periodically modulated
in one, two or three spatial dimensions [5]. For example, the
simplest 1D-PC structure consists of dielectric thin film layers
with a periodic alternating high and low refractive index (RI) to
form a distributed Bragg reflector (DBR). The coherent effects
of scattering and interference result in a dramatic modification of
the dispersion relation for light traveling in the DBR [6]. Under
proper conditions, a DBR can sustain surface wave, named
Bloch surface wave (BSW), confining the electromagnetic field
at the interface between the DBR and the medium under examination [7], [8]. Such evanescent wave is typically used to
probe biomarker recognition events at the surface in biosensing
applications [9].
Another class of 1D-PC are those based on 1D grating structures and quasi-guided or leaky modes named GMRs (guided
mode resonances) [10], [11]. Analogously to typical guided
modes, GMRs localize the energy in the slab and can be efficiently coupled into or out of the slab facilitating the (bio)sensing
features. In contrast, 2D and 3D photonic crystals can be designed using different types of symmetries with implications
for the localization of the electromagnetic (EM) field in such
structures. The most common configuration of 2D-PC is a
silicon or gallium arsenide slab in which air holes are etched
periodically resulting in a highly ordered distribution of the RI
arranged in a planar dielectric waveguide [12]. As a consequence
of such arrangement, the light is guided in the slab whereas the
confinement in the third dimension is guaranteed by the total
internal reflection. PC cavities are obtained by inserting defects
in the periodic slab at defined locations confining or the guiding
specific EM modes in order to maximize the interaction with the
medium under examination [13].
Recently, an emerging class of 3D-PC based biosensors is
overlooking on the panorama of biosensing applications.
Typically, they are constituted by direct opal (dO), crystalline
colloidal array (CCA) and inverse opal (IO) structures. Due to
3D ordered nanostructures, dO, CCA and IO based PCs can be
exploited in biosensor applications [14]–[16]. Opal structures
consist of dielectric spheres arranged in a fcc lattice while
artificial opals or CCA can be prepared by self-assembling
dielectric spheres in a colloidal solution. On the other hand, IO
structures with their porous and interconnected architecture span
several technological applications from optoelectronics, passing
for the energy storage to sensor and biological applications
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JOURNAL OF LIGHTWAVE TECHNOLOGY, VOL. 39, NO. 12, JUNE 15, 2021
[16], [17]. Recently, there has been an increasing interest in
the integration of photonic crystals with optical fibers (PCF).
Such PC integration consisted of the structuring of the fiber
tip with a grating [18] or, more recently, of a D-shaped fiber
with modified geometry [19]. In both cases, the figure of merit
and sensitivity are promising for biosensing applications. Nevertheless, at the moment, there are no experimental evidences of
cancer biomarker detection using such devices.
The use of PC as a diagnostic tool may introduce advantages
in terms of device miniaturization (μ-TAS), multiplexing, fluidic
design and integration allowing to process low amount of samples from patients. This permits to make complex diagnostic
analyses more efficient with lower cost, energy, and chemical consumption than conventional systems and dramatically
improve performances of biosensors in terms of sensitivity,
accuracy and limit of detection (LoD) [20].
In this context, due to their critical role at all stages of disease,
cancer biomarkers have gained more and more attention in recent
years. In particular, according to the US National Cancer Institute (NCI) a cancer biomarker is “a biological molecule found
in blood, other body fluids, or tissues that is a sign of a normal or
abnormal process, or of a condition or disease,” such as cancer
[21]. Therefore, cancer biomarkers include circulating tumor
DNA (ctDNA) [22], [23], miRNA (and other non-coding RNA)
[24], proteins [25], enzyme [26], metabolites [27], as well as
circulating tumor cells (CTCs) [28]. Such a tremendous variety
of biomarkers are produced either by the tumor itself or by other
tissues, in response to the presence of cancer or other associated
conditions, such as inflammation. Indeed, they are generally
detected in the circulation (whole blood, serum or plasma) or
in secretions and excretions (saliva, urine, stool) allowing for
easy, non-invasive and serial approaches to diagnosis, or can
be tissue-derived, requiring either biopsy or special imaging
evaluation. However, in order to be incorporated into routine
clinical care, they have to undergo a rigorous analytical and
clinical validation guaranteeing a real clinical utility.
This review wants to give an overview on photonic crystalsbased biosensors for cancer biomarker detection. The present
review wants to describe the basics of the most common used
1D, 2D, and 3D-PC configurations providing an overview of the
main biosensing applications involved in cancer diagnostics and
therapy monitoring.
II. 1D-PCS FOR CANCER BIOMARKER DETECTION
A. Basics of 1D-PC Based Biosensors
A DBR structure consists of a periodic alternating high and
low RI dielectric thin film layers to form a dielectric mirror.
Indeed, multiple reflections from consecutive layers provide
constructing interference producing a total reflection. The understanding of such optical phenomenon permits to treat theoretically 1D-PC as DBR. More precisely, the optical thicknesses
of these dielectric layers are designed to match the following
equation (Eq. 1):
dH nH = dL nL =
λ
4
(1)
Fig. 1. (a) DBR geometry, TE polarized BSW with its evanescent field (red
profile); (b) simplified read-out configuration; (c) calculated angular reflectance
at λ1. The double green arrow marks the label-free angular operating window;
(d) radiant intensity calculated for isotropically oriented dye molecules (Dyomics 647) placed at the surface of the DBR. The double red arrow marks the
fluorescence angular operating window. Reproduced with permission from [32].
where dH nH and dL nL are the optical thicknesses for the high
and low RI layers, respectively. Generally, the fabrication of such
multilayers is obtained by RF/magnetron sputtering or plasma
assisted evaporation process under high vacuum conditions.
Typical materials used in the DBR fabrication are SiO2 , Ta2 O5
and TiO2 [29].
The optical working principle is based on multiple reflections
from consecutive layers providing constructive interference and
result in total reflection. The reflected wavelengths that reside in
a PBG region cannot propagate at normal angle of incidence. A
pioneer work in 1978 theoretically [30] and experimentally [31]
demonstrated that a special type of surface wave can appear on
a truncation surface of semi-infinite DBR multilayer (Fig. 1(a)).
These surface states, known as Bloch surface waves (BSWs),
are spatially confined electromagnetic waves that exploit the
PBG of the photonic crystal to guide a surface evanescent wave
(Fig. 1(a)).
The penetration of evanescent wave in the liquid under investigation can be in principle engineered as a function of the
thicknesses and RIs of the multilayer. In operative cases, the
penetration length of the BSW is in the order of 100–200 nm.
The use of dielectric materials with low optical losses results in a
propagation of the BSW over longer distance. In addition, DBR
can be designed in order to sustain both s-polarized (electric
field parallel to the surface, TE) and p-polarized (electric field
perpendicular to the surface, TM) BSWs and can be tuned to
operate at almost any wavelength varying the geometry and/or
the dielectric materials of the multilayer. As shown in Fig. 1(b),
the label-free operation mode can be obtained exciting a BSW
at λ1 by means of a prism coupler in the Kretschmann-Raether
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SINIBALDI: CANCER BIOMARKER DETECTION WITH PHOTONIC CRYSTALS-BASED BIOSENSORS: AN OVERVIEW
Fig. 2. (a) SEM image showing the surface structure of the 1D-PC slab;
(b) schematics of the periodic surface structure fabricated in a low RI silicon
dioxide (SiO2 ) layer that is overcoated with a high RI film of TiO2 . Reproduced
with permission from [40]. (c) schematic diagram of the detection instrument
and a typical reflected intensity as a function of the wavelength. Reproduced
and adapted with permission from [41].
configuration [33], [34] producing a dip in the angular reflectance. As shown in Fig. 1(c), by tracking the angular position
(θRES ) of the TE dip, it is possible to monitor changes in the RI
as well as molecular interactions occurred at the surface. Again,
through the prism, a resonant excitation of a dye molecules
anchored at the surface can be obtained at the maximum of
the dye absorption spectrum (λ2 ). As depicted in Fig. 1(d),
the angular emission of the dye molecules in the presence of
the DBR is modified and re-directed in the collecting system
appearing as two replicas of the dye emission spectrum decoupled by means of TE and TM BSWs, respectively. Such an
enhanced fluorescence mode offers a sharp improvement of the
resolution of the technique and makes it particularly attractive
for biosensing applications [35], [36].
Another common 1D-PC geometry used in biosensing applications exploits periodically modulated thin film and GMRs.
GMRs are quasi-guided or leaky modes that are able to readily
couple the energy confined in the 1D-PC slab to the external
radiation. This feature allows for an efficient way for coupling
power into and out the slab facilitating the sensing mechanism.
Typically, this class of biosensors is obtained using a lowcost nano-replica molding manufacturing approach or thermal
growth to define a periodical low RI grating. Then a high RI
coating layer of TiO2 is deposited by RF sputtering onto the
nanostructured grating as shown in Fig. 2(a) and 2(b) [37]. The
RI modulation acts as a wavelength-scale grating that permits,
under proper conditions, to excite the GMR. Moreover, the high
RI coating behaves as a non-homogeneous waveguiding layer
where the guided mode scatters at each optical discontinuity
giving rise to coherent scattering.
By manipulating the grating period (Λ), material RIs, angle
of incidence and polarization, the phase can be tuned ensuring
destructive interference in transmission between the forward
diffracted zeroth order (for normal incidence) and the light
scattered upward by the leaky mode. As shown in Fig. 1(c),
this condition of phase matching produces a sharp peak in the
reflection spectrum with typical Q factor in the range of 100–200
[12]. The resonance wavelength peak is defined in Eq. 2 taking
into account the Λ and the effective RI (nef f ) at resonance
condition as follows:
λRES = nef f Λ
(2)
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Based on this resonance principle, the central wavelength
λRES of the GMR structure can be tuned by altering the angle of
incidence. The resonant features of such a 1D-PC slab make it
highly sensitive to changes in the dielectric permittivity close to
the slab surface providing a suitable tool for biosensing applications in label-free, enhanced fluorescence and spatially resolved
imaging. The label-free sensing mechanism foresees a shift of
the resonance wavelength peak due to molecular binding which
is proportional to the amount of molecules bound at the 1D-PC
surface. Likewise, the exploitation of 1D-PC slabs to obtain
large field enhancement factor in presence of fluorophore-tagged
proteins has been widely demonstrated [38]. Indeed, in this case,
the 1D-PC surfaces are engineered to provide a resonance peak
at the excitation wavelength of the fluorescent dye. This provides
an enhanced excitation in the evanescent region of 100–200 nm
above the 1D-PC surface. At the same time, a second resonance
peak (generally obtained with a different polarization) close
to the emission wavelength of the dye provides an enhanced
fluorescence extraction. The two latter effects are multiplicative
and can give rise to an overall signal enhancement of about 600
with respect to a glass substrate [39].
B. 1D-PC: Applications to Cancer Biomarkers Detection
As mentioned above, a SiO2 /Ta2 O5 -based DBR sustaining
BSW can combine label-free and fluorescence detection modes.
One of the most remarkable application of such real-time optical
technique is the detection of clinically relevant concentrations
of HER2 in cell lysates. Indeed, HER2 (epidermal growth factor receptors) is a targetable oncogenic driver associated with
aggressive breast cancer subtypes. Typically, the DBR surface
can be tailored with three different probe antibodies to form a
capturing array as a starting point for a sandwich assay. As an
example, a capturing antibody array can be formed by HER2
specific (Anti-HER2, S) and HER2 non-specific antibodies
(Anti-MHC I, R1; Anti-MHC II, R2). Once the DBR surface
is biologically active, a blocking step in bovine serum albumin
(BSA) is introduced to properly passivate the surface of the
biosensor.
In Fig. 3(a) are reported the sensograms for the three probing
spots as a function of the assay progression. Two injections
characterize the assay: first, a HER2 positive/negative cell lysate
is allowed to react with the DBR biochip; second, an Anti-HER2
antibody (conjugated with Alexa Fluor 647) solution is introduced to amplify the label-free signal as well as specifically
introduce a fluorescent tag. In this case dilutions of HER2
positive (SKBR-3)/negative (Colo38) cell lysate were measured
in both operating modes. The label-free signal as a function
of HER2 concentration in a SKBR-3 positive cell lysate were
recorded showing a clear binding and an increasing angular
shifts.
The latter results were confirmed by the integrated fluorescence intensity results associated to the same experimental
determinations. In Fig. 3(b) were displayed all the spots for all
dilutions of the SKBR-3 and Colo38 demonstrating an increase
of the fluorescence signal for SKBR-3. The estimated LoD in
label-free was set to 17 ng/mL whereas, for the fluorescence
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JOURNAL OF LIGHTWAVE TECHNOLOGY, VOL. 39, NO. 12, JUNE 15, 2021
Fig. 3. (a) Label-free signal recorded during a complete HER2 bio recognition
assay in a SK-BR-3 cell lysate (sample P1). The three sensograms correspond
to the specific (S, black) and non-specific (R1 red and R2 blue) antibody spots;
(b) Background subtracted power W after exposition to a second Anti-HER2
labeled with Alexa Fluor 647 and for different SK-BR-3 and Colo 38 samples.
The different colors denote fluorescence values obtained for the three sensitive
spots on the 1D-PC biochips. Reproduced with permission from [9]. (c) Picture of the plastic 1D-PC biochip with integrated micro-optics and polymeric
microfluidic cover; (d) Fluorescence emission recorded by the imaging system
at the end of an assay, where Ang2 was detected; (e) Angular emission profiles
along the dashed lines for S and R spots. Reproduced with permission from [43]
© The Optical Society.
operation mode, the DBR attained a LoD of 0.3 ng/mL for HER2
in cell lysates. Such latter resolution meets the international FDA
guidelines and recommendations (15 ng/mL) for diagnostic
HER2 assays [9].
Proangiogenic and angiogenic factors such as Angiopoietin-2
(Ang2) and vascular endothelial growth factor (VEGF) have
been widely investigated due to their function in tumor vasculature and in other pathological conditions associated with
endothelial dysfunction [42]. The biochips used in these works
were fabricated by depositing purposely designed DBR directly
on molded plastic substrates (TOPAS 6013). Such approach
led to a disposable and low-cost biosensor units, which are
advantageous to the practical applications. As shown in Fig. 3(c),
the integrated 1D-PC substrate was equipped with a two components injection molded flow cell (microfluidic cover). By means
of the same technique, and using DBR-based biochips, Ang-2
and VEGF were detected in human plasma and cell lysates,
respectively. In Fig. 3(d), a typical fluorescence output with
signal (S) and reference (R) spots has been reported. In Fig. 3(e)
are reported the angular emission spectra for both a S and R
regions. The LoDs attained in enhanced fluorescence mode for
Ang-2 and VEGF were 6 and 0.65 ng/mL, respectively [43],
[44].
A similar approach with a different DBR geometry was
reported for ovarian and breast cancer biomarkers detection.
Also in this case, the DBR was a SiO2 /Ta2 O5 -based multilayer
in a real-time flow detection. This label-free high-precision
biosensing technique allows monitoring of molecular and cellular interactions using independent recording of the total internal
reflection angle and the excitation angle of the BSW surface
wave. Such a technique permitted to simultaneous detect the
Fig. 4. (a) Transmission spectrum (T) of the 1D-PC at normal incidence. The
resonance wavelengths for the TM (solid curve) and TE (dashed curve) polarizations occur at 629 and 690 nm, respectively; (b) Fluorescence comparison of
on and off-resonance measurements for different functional assays in the array.
Reproduced with permission from [38]; (c) Bright field and (d) PCEM image of
Panc-1 cells attached to the 1D-PC surface; (e) Representative regions of cellular
attachment. Selected areas of the PCEM image from beneath a cell show the
peak shift of a typical Panc-1 cell (about 1.0 nm). Reproduced with permission
from [39].
ovarian cancer biomarker CA-125 and two breast cancer markers, HER2 and CA-15-3. Although, the biosensing was obtained
in buffer solutions (PBS 1X) and with non-homogeneous interaction times, the LoD estimated for HER2 was 0.62 pg/mL with
a linearity range of up to 50 pg/mL [45].
GMR-based on 1D grating were also used for the detection of
several cancer biomarkers in serum samples [40], [46]. In this
case, the discrimination between different subtypes of breast
cancer and benign controls has been successfully obtained [38].
In such an application, the 1D grating was obtained starting from
a polymeric material (UVCP) on plastic substrates by means of
SiO2 and TiO2 coating procedures. In this sandwiched structure,
the presence of TiO2 was necessary to establish the localized
GMR. Such a device allowed the analysis of a panel of 24 cancer biomarkers using commercially available confocal microarray scanners. The enhanced fluorescence excitation/extraction
mechanism is based on the simultaneous presence of both a TM
and a TE modes by illuminating at normal incidence and using
a Cy5 labeling system.
Under a broadband illumination, the 1D-PC system was designed to have a TM resonance at the Cy5 excitation wavelength
and a TE resonance at the Cy5 emission for enhanced extraction
as shown in Fig. 4(a). Once functionalized with an epoxysilanebased reagent, the arrays onto the 1D-PC surface were incubated
with different capture antibodies and then blocked with casein.
To generate standard curves the 1D-PC slides were incubated
overnight with a mixture of antigens and 0.1% of casein. The
sandwich assays were completed by means of biotinylated antibodies and Cy5-conjugate streptavidin. In Fig. 4(b) (left) it
is reported a typical fluorescence on-resonance array output.
As an example, in Fig. 4(b) it is reported the fluorescence
intensity distribution off and on-resonance for 13 of the 24
cancer biomarkers investigated. The LoDs obtained with such a
1D-PC in on-resonance conditions ranged from 1.9 pg/mL for
tumor necrosis factor alpha (TNF-α) to 1.9 ng/mL for urokinase
plasminogen activator receptor (uPAR) [38].
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SINIBALDI: CANCER BIOMARKER DETECTION WITH PHOTONIC CRYSTALS-BASED BIOSENSORS: AN OVERVIEW
A similar 1D-PC implementation was used also to monitor
miRNA. In particular, the target miRNA sequence miR-21,
a sequence implicated in the progression of breast cancer,
was detected reaching a LoD of 0.60 pM [47]. Another dual
polarization GMR-based 1D-PC was used for the label-free
detection of cancer biomarkers. In this experiments, the TE
and TM resonance peaks were tracked and used to distinguish
any physical perturbation or fluctuation from binding events
in a self-referencing configuration. In particular, TNF-α and
calreticulin cancer biomarkers were detected obtaining a LoD
of 156 and 390 ng/mL, respectively [48].
By using the same strategy, time-resolved label-free imaging
of cell-surface interactions with submicron resolution has been
also performed. In this case a photonic crystal enhanced microscopy (PCEM) was used to measure the resonant wavelength
peak value on a pixel-by-pixel basis over the 1D-PC surface.
By evaluating the shift in wavelength an image of cell attachment density was recorded in aqueous media [39]. This
permitted to study, without the use of dyes or stains, subtle
variations in cell adhesion strength within a single cell. As cell
attach and spread, positive wavelength shifts are recorded due
to an increase of the concentration of cellular material within
the evanescent tail of the EM. As an example, in Fig. 4(d),
is reported a PCEM image for human pancreatic cancer cells
(Panc-1) compared to a bright field image (Fig. 4(c)). In Fig. 4(e),
representative spectra are shown from inside and outside the
cell region demonstrating to provide information about the geometry of the attachment with significant implications for both
classification and metastatic potential of tumor cells. Moreover,
sub-cellular variations in the resonant peak are also indicative
of a gradient in the strength of attachment suggesting formation
of lamellipodia in higher concentration regions [39].
III. 2D-PCS FOR CANCER BIOMARKER DETECTION
A. Basics of 2D-PC Based Biosensors
The most common 2D-PC configurations used in biosensing
are two-dimensionally periodic slabs consisting of triangular or
square lattice of air holes or pillars/rods. Such 2D periodic slabs
can be easily defined and fabricated in photoresist and transferred into high RI materials such as silicon, gallium arsenide or
silicon nitride using standard dry-etching techniques [49]. The
most common substrate used for 2D-PC based optical devices
is silicon-on-insulator (SOI) where a thin layer of Si (<400 nm)
is placed on top of a SiO2 thicker layer. As an example, in
Fig. 5(a) and 5(b), are reported different 2D periodic structures
obtained on SOI. The periodic arrangement of holes/pillars on
the substrate produces a modulation of the RI embedded in a
planar dielectric waveguide where the light is guided in-plane
by the PBG while out-of-plane (third dimension) by TIR. The
presence of a PBG permits to confine EM modes at defined
locations or to guide them along pre-established path with the
aim to favor the interaction with the liquid under investigation.
The introduction of defects in the periodic arrangement of RI
can be obtained by removing holes/pillars or replacing them with
different element geometries. The final result is the formation
of waveguides (PCWs) and microcavities (PCMs) or, by virtue
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Fig. 5. (a) SEM image of the L3 PCM located 2 rows away from a W1
PCW. Reproduced with permission from [50]; (b) SEM image of the 2D hollow
photonic crystal cavity for optical trapping and Gram-type differentiation of
bacteria. Reproduced with permission from [51]. (c) SEM image of a heterostructure microcavity of pillar-array. Reproduced with permission from [52] © The
Optical Society. (d) typical L3 PCM side coupled W1 scheme; (d) typical hollow
photonic crystal cavity; (f) typical PCM heterostructure scheme. In the insets
are reported the dip/peak positions of the transmission spectra when RI of the
liquid under investigation is changed.
of the exceptionally small element sizes, nanocavities [13,] with
available states in the PBG characterized by a high-degree of EM
localization. PMCs can be fabricated by inserting point and/or
line defects or using heterostructures. On one hand, the removal
of groups of periodic elements or the introduction of elements
with different geometry provides a stronger localization of the
light in defined areas of the slab. On the other hand, heterostructures are usually obtained altering the period of the 2D-PC in a
specific section as shown in Fig. 5(c) and 5(f).
As sketched in Fig. 5(d) and 5(e), PCMs are characterized
by a dip in transmission spectra corresponding to the resonant
wavelength of the mode (λRES ). Moreover, when the PCM is
created using a heterostructure, a peak can be found in the
transmission spectra at λRES (Fig. 5(f)). The localization of
the light in PCMs can be described by the definition of two
characteristic parameters: the quality factor Q and the mode
volume Vm of the cavity. The quality factor is usually expressed
as in 3:
WE
λRES
= ωRES
(3)
Q =
δλ
PR
where Δλ is the bandwidth of the resonance, ωRES is the
resonant frequency, WE is the energy stored inside the cavity
and PR is the power dissipated out from the cavity. The mode
volume Vm is the space where the electric field is confined and
resonantly enhanced, the order of magnitude is provided by 4:
3
λ
(4)
Vm ∼
n
Typical values for PCMs are Q ∼ 102 − 106 for Si (reaching
Q > 107 for nanobeam cavities) and Q > 106 for GaAs while
for Vm ∼ 0.39 − 0.71 (λ/n)3 [53], [54].
One of the most common transmission drop-resonance architecture in passive hexagonal silicon-based 2D-PC is sketched
in Fig. 5(d). A linear PCM coupled to a photonic crystal waveguide (PCW) were obtained removing 3 adjacent (L3) and one
complete row of holes (W1), respectively [13]. The change in
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Fig. 6. (a) SEM image of the side-coupled L13 PCM; (b) Multiplexed simultaneous specific detection of ZEB1 in lung cancer cell lysates in the four arms (A1)
BSA, (A2) mouse IgG1, (A3) anti-ZEB1 and (A4) anti-MYC 9E10 antibodies.
Baseline spectra for each arm is shown in black. Resonance wavelength positions shown in red, blue, green, respectively with sequential addition of 3-day
induced lung cancer cell lysates, mouse IgG1 and anti- MYC 9E10, respectively.
Reproduced with permission from [61]. (c) SEM image of nanobeam cavity and
photonic circuits consisting of waveguides (green lines), cavities (in red boxes)
and grating couplers (in black boxes), (d) residual resonance shifts for increasing
concentration of CEA. Adapted with permission from [62] © The Optical
Society. (e) SEM image of the L13 PCM with nanoholes; (f) Resonance shifts
for L13 and L13 with nanoholes for different concentrations, the dashed line
represents the spectrum analyzer detection limit. Reproduced with permission
from [63].
the RI in the external medium generates a shift in the resonant
dip Δλ as shown in the insets of Fig. 5(d). Similarly, when a
biomolecule absorbs at the PCM surface, an alteration of the
total Q is registered providing a shift in wavelength. Typical
outputs of such a label-free detection mechanism [55], i.e., when
a change in the RI is applied, are also sketched in the insets of
Fig. 5(e) and 5(f).
In similar 2D-PC devices, the control of the speed of the
light represented a powerful tool in biosensing applications.
In particular, the slow-light phenomenon can be successfully
exploited in PCW and PCM biosensors due to the enhanced
light-matter interaction [56], [57]. The combination of slowlight engineering with side-coupled PCMs can lead to higher
sensitivities and Q factors introducing longer cavities [58]. The
lengthening of the cavity (waveguide) produces an increase of
the Q factor moving the modes close to the band edge of the
waveguide. This corresponds to a group index increase that
enhances the slow-light effect [59], [60].
These features can lead to multiplexed and error-corrected
biosensing configurations with the possibility to create photonic
circuits. Such systems exploit the well-known semiconductor
fabrication processes pushing the integration and the scalability
of the devices from micron to nanoscale.
B. 2D-PC: Applications to Cancer Biomarkers Detection
This section reports on different in-plane 2D-PC structures
to quantify various cancer biomarkers. In particular, one of
the most commonly structure used in biosensing foresees a
PCM coupled to a PCW in Si on a SOI substrate. As shown
in Fig. 6(a), the PCW was a W1 line defect waveguide on a
slab with a uniform distribution of holes (lattice
√ constant a =
400 nm) and with a standard width of w = 3 a. The silicon
slab thickness and air holes diameter were h = 0.58a and
d = 0.57a, respectively. The linear L13 PCM was produced
positioning 13 missing holes two period away from PCW [61].
For this configuration, a Q factor of approximately 1.3 104
was confirmed experimentally. Moreover, the device was implemented to provide multiplexed resonance cavities arranging in
parallel 4 detection arms with one side-coupled PCM each. This
permitted to have simultaneous detection obtaining duplicate or
triplicate analyses in the same measurement improving statistics
(see the inset in Fig. 6(a)). In order to detect the biomarker
ZEB1 in NCI-H358 lung cancer cell lysates, the four arms of
the array were immobilized with two specific probe for ZEB1
(anti-ZEB1 and anti-MYC 9E10 antibodies) and two reference
molecules (BSA and mouse IgG1). In Fig. 6(b) were reported the
typical transmission spectra of the side-coupled PCMs device
in PBS buffer in the four detection arms (black spectra). After
interaction with the cell lysate containing native and induced
ZEB complex, a clear shift can be appreciated only on anti-ZEB1
and anti-MYC arms. Besides, no residual shifts were recorded in
the reference arms confirming the specificity of such multiplexed
side-coupled PCM biosensor to ZEB1 cancer biomarker. Such
a biosensor demonstrated the capability to specifically detect
ZEB1 in 2 cells per μL of induced NCI-H358 cells [61].
Similarly, by using the same 2D-PC structure, a complex study
on association and dissociation kinetics were conducted for
interleukin (IL)-10 [60]. IL-10 is a cytokine and appears to have
considerable importance in the development of human gastric
cancer and its immune escape [64]. In this case the authors
attained experimentally a sensitivity of 98 ag (8.9 pg/mm2 ) for
a dissociation constant of 10−10 M with the conjugate pair of
human IL-10 and rat anti-human IL-10 [60].
Side-coupled PCMs were experimentally characterized introducing periodic nano-defects within the cavity. In Fig. 6(e)
is reported a side-couple L13 PCM structure fabricated by
photolithography process and modified with nanoholes defects.
This structure were compared with a conventional L13 PCM (Q
= 1.2 104 ) with the aim to study the optical characteristics in
terms of Q factor and sensitivity in the detection of pancreatic
cancer biomarkers [63]. The PCM were fabricated in Si on a SOI
wafer where the photonic crystal was defined by a triangular
lattice of air holes with hole radius R = 112.5 nm. The radius
the nanoholes was r = 0.4R and were arranged in the center of
the L13 PCM. As a consequence of such a nano-structuration,
an additional sharp resonance with a Q factor of 2.2 104 was
achieved for the L13 PCM with nanoholes in water ambient.
The 2D Si slab was functionalized by means of APTES and
glutaraldehyde prior the immobilization of the probe antibodies.
Three pancreatic cancer biomarkers HGF, MIP1, FAS ligand
were tested to demonstrate the specificity of the L13 PCMs with
and without nanoholes. In Fig. 6(f) were plotted the resonance
wavelength shifts for PCMs L13 and L13 with nanoholes in
different dilutions. Thus the conventional L13 PCM was easily
able to detect the HGF and MIP1 with a LoD of 9.813 pg/mL and
15.437 pg/mL, respectively; while for FAS ligand an ultimate
LoD of 0.334 pg/mL (8.8 fM) was reached thanks to the L13
PCM with nanoholes [63].
In recent years, photonic crystal nanobeam cavities were
largely used in biosensing applications and in particular in cancer
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SINIBALDI: CANCER BIOMARKER DETECTION WITH PHOTONIC CRYSTALS-BASED BIOSENSORS: AN OVERVIEW
biomarker detection. Such photonic systems take advantage
of their ultra-small optical mode volume thus higher Q factor
(order of 104 ) [62]. As shown in Fig. 6(c), this nanobeam
cavity consisted of a perforated Si waveguide resting on a SiO2
substrate with the possibility to produce chips in large quantities
with scalable photolithography technology. The perforation is
patterned in a periodic way to tightly confine the light into the
center of the array via optical interference [62]. The position of
resonance of the nanocavity λRES is perturbed when proteins
bind to the surface of the cavity increasing the refractive index
of the solution under investigation. The excursion of resonance
shift scales as a function of the quantity of proteins bound on
the nanocavity surface and inversely scales with the size of the
optical mode volume. In order to specifically detect carcinoembryonic antigen (CEA), a tumor biomarker that is used to monitor
the colon cancer treatment, the nanocavity was functionalized by
means APTES and glutaraldehyde prior the anti-CEA antibody
bioconjugation. After the anti-CEA tailoring of the nanocavity,
the biosensor was exposed to increasing concentration of CEA
(0.1, 1, 10, 100 pg/mL; 1, 10, 100 ng/mL and 1, 10 μg/mL
in PBS) as shown in the label-free sensogram of Fig. 6(d).
Although, a clear binding signal was observed starting from
10 pg/mL, the LoD was estimated fitting the experimental curve
with Langmuir equation obtaining a LoD of 14 ng/mL [62].
Another example of a photonic circuit on SOI, consisting
of a series of PCMs, was obtained by modifying periodically
the radius of a single air hole neighboring the W1 in order to
introduce defect states in the PBG of the 2D-PC [65]. Also
in this case, the light is transmitted at all frequency except at
the PCM resonance wavelength where a dip is observed in the
transmission spectrum. The presence of multiple PCMs in series
gives rise to multiple dips in the transmission spectrum as each
PCM possesses a unique resonant wavelength. The particular
arrangement of the 3 PCMs in series provides multiplexing
introducing error-correcting mechanisms during biosensing. For
optical detection, the anti-hIgG antibody modified biosensor
were tested with increasing concentrations of target hIgG solutions. The device sensitivity was found to be 2.3 ± 0.24
105 nm/M with an achievable lowest detection limit of 1.5 fg
for human IgG molecules [65].
IV. 3D-PCS FOR CANCER BIOMARKER DETECTION
A. Basics of 3D-PC Based Biosensors
In the last two decades, the improvements obtained in the
fabrication of artificial 3D-PC allowed to explore new configurations for optical biosensing. The most common 3D arrangements used for biosensing consist in the nano/micro
structuring techniques by means of colloidal suspension of
silica (SiO2 ) or polystyrene (PS) or polymethyl-methacrylate
(PMMA) nano/microspheres. This fabrication technique is one
of the most favoured and low cost methods for the formation of
3D-PCs as artificial opals [66], [67]. Indeed, a direct opal (dO)
can be produced starting from a colloidal solution through a
self-assembling process to obtain a face centered-cubic (fcc)
[111] lattice of SiO2 or PS or PMMA (inset Fig. 7(a)). A
standard fabrication protocol forecasts: gravity or centrifugal
3877
Fig. 7. (a) Reflectance spectra (R) of the dO and IO nanostructures. In the
inset SEM images for a dO and IO structures [73]; (b) reflectance peak evolution during functionalization and bio-conjugation processes. The inset shows
reflectance spectra for such steps; (c) Optical images showing contact angles
of a water drop on the dO (top panel) and IO (bottom panel) nanostructures.
Reprinted with permission from [73], [74]. (d) Schematic illustration of dye
solution on PDMS and on 3D-PCs surface under excitation (upper part). The
photographs of dye solution on bare PDMS and on PCs surface under excitation
with a UV lamp (bottom part). Reprinted with permission from [75].
sedimentation, self-assembly, evaporation, dip coating or electrophoresis methods [68], [69]. On the other hand, inverse opal
(IO) can be prepared by filling the SiO2 / PS /PMMA template
by a high refractive-index target material usually Si, TiO2 or
ZnO. After filling the dO structure, a typical IO fabrication
protocol foresees the removal of colloidal crystal templates by
calcination or etching of the dO template (inset Fig. 7(a)). Thus,
an IO 3D-PC consists of closely packed spherical air voids
arranged in a fcc manner and embedded in a scaffold of solid
target material. However, IO can be also obtained by hydrogel
or elastic polymer matrices starting from a PS template [70]
which could be stretched or shrunk in order to tune the initial
intra-elements distance and thus changing the optical response
[71]. Likewise, with 1D and 2D photonic crystals, the periodic
spatial variation of RI leads to the formation of spectral PBG or
photonic stop gaps (PSG). The propagation of the EM waves is
thus forbidden along all directions in the case of PBG or along
specific directions for PSG.
As shown in Fig. 7(a), optical (bio)sensing can be obtained by
detecting modification of the PBG or PSG spectral signatures in
the dO or IO reflection or transmission spectra [73]. In Fig. 7(a)
are reported the spectral signatures as reflection peaks for dO
and IO 3D-PCs. It is clear from Fig. 7(c) the difference in the
sensing mechanism for the two structures. In the case of dO, the
change of refractive index of the surrounding medium provides
a blu/red shift of the peak position while, for an IO, the change
of the peak position is mainly due to infiltration of liquid under
examination within the structure. Such sensing features can be
summarized in Fig. 7(c) where contact angle measurements are
reported for the dO and IO structures. Moreover, the desired peak
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JOURNAL OF LIGHTWAVE TECHNOLOGY, VOL. 39, NO. 12, JUNE 15, 2021
position can be easily tuned by varying the sphere diameter for
both dO and IO structures. The larger is the sphere/void the
larger is the sensitivity [73]. For the sake of clarity, it is helpful
to introduce the analytical expression for the spectral position
of the fundamental PSG of a 3D-PC opal structures.
In the direction orthogonal to the [111] crystal plane, the PSG
peak wavelength λP can be retrieved in 5 by the Bragg condition:
λP = 1.633 d [fs n2s + (1 − fs ) n2r ]
(5)
assuming that the average RI is determined from sphere filling
fractions (fS ) and from RI of constituent materials (nS and
nr ). The parameter d represents the diameter of the spheres.
In general, closely packed opals show a sphere filling fraction
of 0.74, while the residual voids fill the (1- fS ) = 0.26 volume
fraction.
Indeed, in the case of dO, nr can be changed by infiltration
of the liquid under investigation into the voids while nS is kept
constant. Contrarily, in IO, nS is related to voids and nr ,that
represents the framework, remains with a constant RI during
biosensing measurements. Although the PSG peak wavelength
λP is more sensitive to the RI changes of nS with respect to nr ,
due mainly to the large volume-filling fraction, only recently IO
biosensors have been used in biosensing applications [73].
As described in Fig. 7(b), with such a mechanism it is possible to study molecular interactions at the inner surface of an
IO structure when functionalization/bio-conjugation processes
are performed. For example, to detect influenza A (H1N1)
virus, a SiO2 IO structure has been modified to be specific to
hemagglutinin (HA), a homotrimeric glycoprotein found on the
surface of influenza viruses [74]. The SiO2 surface is allowed to
react with (3-Aminopropyl) trimethoxysilane (APTMS) in order
to guarantee amine groups at the surface. Then, the APTMS
modified SiO2 -IO is conjugated with NHS-PEG4-maleimide
(M), an heterobifunctional crosslinker, with the aim to link the
surface with a thiolated protein G (PtG). The main role of such
a protein is to well orient the antibody (Ab, Anti-HA) specific
for HA. All these functionalization and bio-conjugation steps
can be followed in real-time monitoring the shift of the reflected
peak as shown in Fig. 7(b).
Alternatively to label-free, crystalline colloidal array (CCA)
are also used as biosensors to enhance fluorescence signal coming from specific reactions occurred at the transducer surface.
A typical CCA fabrication process forecasts a vertical dropping
of PS spheres on a PDMS substrate to form a nanostructured
surface. When the excitation wavelength is aligned with the
CCA diffraction peak, the EM fields on the surface of the
nanostructured film are enhanced.
This permits to resonantly excite fluorophores at the surface
guarantying a more efficient delivery of energy to the emitters
with respect of an optically passive structure such as glass (Fig
7(d)). Furthermore, whether the emission wavelengths of the
fluorophore overlap with the CCA diffraction wavelengths, the
emitted light can be redirected toward the detection equipment
enhancing the extraction [71]. Such a peculiar characteristic
is particularly interesting for the development of fluorescence
biosensors with high signal-to-noise ratio (SNR) thus improved
LoD.
Fig. 8. (a) In situ reflectance spectra showing the immobilization of hIgG,
the blocking of BSA, and the binding of goat anti-hIgG on the pore surfaces.
(b) dependence of the red-shift on the concentration of goat anti-hIgG. Inset:
SEM image of the IO TiO2 structure. Reprinted with permission from [72];
(c) comparison of cells imaging on flat glass (site 1 on the left) and IO substrate
(sites 2 and 3 on the right) with CLSM; (d) spectra of the C6 measured by CLSM
from site 1 (on glass), site 2 (on the gap of IO), and site 3 (on IO interface),
respectively. On the right side SEM images of captured cancer cells on flat glass
(G) and on IO nanostructure (size 415 nm). On IO structure the captured CTCs
show an amplification of surface filopodia. Reprinted with permission from [76].
(e) Fluorescent images of let-7a miRNA-initiated RCA reactions at different
let-7a concentrations on PCs surface (upper panel) and on bare PDMS (lower
panel); (f) relationship between the fluorescence intensity and the concentration
of let-7a miRNA for the two cases. Reprinted with permission from [75].
B. 3D-PC: Applications to Cancer Biomarkers Detection
One of the most important features of an IO film is to produce
a high-quality and uniform colloidal crystal. In the case reported
in Fig. 8(a), the authors made use of a vertical lifting method to
fabricate the opal template. In order to obtain a highly ordered
colloidal crystal template, mono-dispersed PMMA spheres with
an average diameter of 275 nm were assembled on a glass
substrate. Then, an aqueous suspension of TiO2 nanoparticles
(15 nm) is infiltrated in the interstitial space of the PMMA
lattice. After calcination, an ordered TiO2 IO film with a distance
pore-to-pore of 210 nm was obtained. To make the TiO2 IO
film biologically specific, the authors adsorbed on the surface of
the micro-cavity a human IgG (hIgG, black curve). After hIgG
incubation, a blocking step with BSA was performed producing
an incremental red shift in the diffraction peak as shown in
Fig. 8(a)(red curve). The bio-conjugated TiO2 IO was then
exposed to increasing concentration of goat Anti-hIgG (0.10,
0.25, 0.5, 1.0, 2.0) mg/mL keeping each solution for 1h at room
temperature and followed by a rinsing step. The shift obtained
by goat Anti-hIgG solutions increased proportionally with their
concentrations reaching a plateau in the calibration curve (inset
in Fig. 8(b)). The biosensing characteristics can be evaluate from
the LoD in about 1 μg/mL that corresponds to a 1.5 pg/mm2
according to the internal surface area of the pores. Several tricks
can be used to improve i.e., lower the LoD in IO biosensors.
For example, one can increase the size of the pores influencing
directly the sensitivity [72] or increase of the density of the
capturing elements immobilized on the pore surface.
TiO2 IO were successfully used also for a highly-efficient capture of circulating tumor cells (CTCs). Such a CTCs immunocapture was obtained by a combination of magnetic nanoparticles conjugated with coumarin 6 (C6) dye and antibodies specific
for epithelial cell adhesion molecules (anti-EpCAM). Also in
this case, the TiO2 IO was formed starting from a PMMA opal
template made of sacrificial 415 nm diameter spheres. In the
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SINIBALDI: CANCER BIOMARKER DETECTION WITH PHOTONIC CRYSTALS-BASED BIOSENSORS: AN OVERVIEW
CTC capture experiments, MCF-7 cancer cells labeled with
coumarin 6 through the staining process with nanoparticles were
imaged with a confocal microscopy system (CLSM) to evaluate
the capturing efficiency on the IO surface.
As shown in Fig. 8(c), IO (right) and glass (left) substrates
exhibited vastly different capabilities in catching MCF-7 cancer
cells [76]. The spectra of C6 from cells on glass (G) and on IO
were recorded by CLSM and displayed in Fig. 8(d). C6 spectra
collected from the three regions, glass substrate (1), glass gap
(2) and IO interface (3) revealed a more intense fluorescence
signal from the IO interface demonstrating improved catching
features for the TiO2 IO. The enhancement in MCF-7 capturing
of TiO2 IO interface is evidenced in Fig. 8(d) (right) where a
larger number of filopodia (40-46) per cell were registered with
respect to those established on glass (16-21) per cell [76].
Three dimensional ordered macro-porous polymeric materials were also used as biosensors for hIgG detection. In this case,
a colloidal crystal made of 190 nm SiO2 spheres was used as a
template. After infiltration of solution of polystyrene in toluene
and a treatment with 4% of hydrofluoric acid a polystyrene film
was obtained in a free-standing form (IO). As already described
in previous reported cases, the diffraction peak experiences a
red shift when immobilization with goat anti-hIgG and passivation were performed. Such a shift can be ascribed to a double
contribution: the change in the refractive index of the protein
solution and the adsorption of the proteins on the pore surfaces.
The polystyrene IO showed a concentration-dependent binding
from 0.01 to 2.5 mg/mL with a sensitivity of 0.01 mg/mL [77]
one order of magnitude better than a TiO2 IO structure reported
below.
Another interesting approach in 3D-PC involving colloidal
crystal PS self-assembled spheres was adopted for cancer
biomarker detection. In this particular case, a serum circulating
miRNA Let-7a, is identified as biomarker for the diagnosis
of non-small cell lung carcinoma (NSCLC) and the detection
was assisted by an enzymatic amplification mechanism (rolling
circle amplification, RCA). Fig. 8(e) shows RCA fluorescent
images at different Let-7a miRNA concentrations on 3D-PC
(upper panel) and PDMS (lower panel) substrates. As shown in
Fig. 8(f), the enhancement provided by the presence of 3D-PC
amplifies fluorescence signal thus the sensitivity of the biosensor
reaching a LoD as low as 0.1 aM. The improvement in the
LoD is of 5 order of magnitude between PDMS and 3D-PC.
3D-PC RCA fluorescent signal was finally used to discriminate
healthy donors from NSCLC patients presenting lower levels
of Let-7a miRNA in serum [75]. However, a similar approach
was applied also to detect miRNA-21 levels in breast cancer
cells such as MCF-7 [78]. More recently, a new kind of 3D-PC
structure has been obtained as a hybrid combination of molecular
imprinted polymer (MIP) and IO for the specific detection of
human serum albumin (HSA). The biosensor responded linearly
to the logarithm of HSA concentration in solution in the range
of 15 to 150 fM [79].
V. CONCLUSION
In this work, a wide range of PC-based biosensing modalities
were comprehensively reviewed. Since their first introduction as
3879
transducers, photonic crystals performed the role of enhancing
element to maximize the light-matter interaction. Due to the
fact that the performance of such biosensors can be exhaustively
defined with their associated LoD, such a parameter was selected
as the main metric in the review. For what concerns the 1D
periodic structures, two different configurations of 1D-PC were
reviewed such as DBRs and 1D gratings exploiting BSW and
GMR modes, respectively. Such 1D-PC-based biosensors were
intensively developed in the last two decades for the detection
of cancer biomarkers adopting both label-free and enhanced
fluorescence mechanisms.
Thanks to these features for DBR biosensors based on BSW
the LoD attained for HER2 ranged from 0.3 ng/mL in complex
biological media down to 0.62 pg/mL in simple buffer solutions.
On the other hand, the 1D grating GMR biosensing were demonstrated in a more heterogeneous range of applications from the
detection of tumor-related proteins passing through miRNA to
entire cancer cells. The LoD was obtained over a large spectrum
of cancer biomarker and spanned from 1.9 pg/mL for TNF-α to
1.9 ng/mL for uPAR.
Alternatively, the use of 2D-PC based biosensors was also
studied and explored in cancer biomarker detection. Starting
from well-consolidated fabrication strategies, the 2D-PCs were
fabricated typically using lithography and etching techniques.
Such a feature permitted to integrate microarrays of sensing
elements with the aim to simultaneously detect tumor-related
proteins using ultra-small sample volumes. The LoD attained
strongly depends on the geometry and on the defects introduced
in the guiding structure. For example, for crystal nanobeam
cavities, that take advantage of their small optical mode volume
thus higher Q factor, the LoD obtained for CEA biomarker was of
14 ng/mL. This latter value was dramatically improved making
use of microcavity structure such as L13 PCM with nanoholes
where FAS ligand was detected reaching an ultimate LoD of
0.33 pg/mL.
Then, a panorama of the most common 3D-PC approaches
used in cancer biomarker detection has been reported. The ease
of their preparation and application (no need of complex optical
set-up) allowed a direct integration of dO, CCA and IO biosensor
elements into photonic or microfluidic chips. In particular, TiO2
IO were successfully used for a highly-efficient capture of CTCs
and for hIgG sensing showing a LoD of 10 μg/mL. By means of
CCA-based photonic crystals, an ultimate LoD of 0.1 aM was
obtained for cancer–related miRNAs using however additional
enzymatic amplification systems.
Moreover, with respect with standard PCF-based biosensing
technique such as fiber Bragg gratings (FBG), Optical fiberbased surface plasmon resonance (OF-SPR) and lossy mode
resonances on D-shaped fibers, flat PC biosensors shows comparable sensing performances in terms of LoD. In particular, in
the detection of HER2 breast cancer biomarker, a taper interferometer cascaded with a FBG and OF-SPR sensors presents
a LoD of 2 ng/mL, and 9.3 ng/mL, respectively [80], [81].
Ultimate biosensing performances in IgG/Anti-IgG interaction
were obtained with lossy mode resonances on D-shaped fibers
reaching LoDs of the order of 0.6 pg/mL [82].
In conclusions, PC-based biosensors hold great potential in
the detection of cancer biomarkers thanks to their ultimate
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JOURNAL OF LIGHTWAVE TECHNOLOGY, VOL. 39, NO. 12, JUNE 15, 2021
performances. In the near future they will be able to provide
a versatile sensing tool to clinical personnel and physicians and
further contribute to medical diagnosis such as health assessment, large scale diagnostic campaign and virus detection.
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Alberto Sinibaldi received the M.Sc. degree in nanotechnology engineering
and the Ph.D. degree in electromagnetism from the Sapienza University of
Rome, Italy, in 2012 and 2016, respectively. From 2012 to 2018, he was a
Research Fellow at Molecular Photonics Laboratory, Sapienza University of
Rome working mainly on optical biosensors for cancer biomarker detection.
Since 2019, he has been an Assistant Professor with the Department of Basic
and Applied Sciences for Engineering with the Sapienza University of Rome.
His research interests include optical biosensors, micro- and nano-optics for
molecular recognition, and functional thin films.
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