A multi-pulse ultrasound technique for imaging of thick-shelled microbubbles demonstrated in vitro and in vivo

Sigrid Berg; Siv Eggen; Kenneth Caidahl; Lars Dähne; Rune Hansen

doi:10.1371/journal.pone.0276292

. 2022 Nov 3;17(11):e0276292. doi: 10.1371/journal.pone.0276292

A multi-pulse ultrasound technique for imaging of thick-shelled microbubbles demonstrated in vitro and in vivo

Sigrid Berg ^1,^*, Siv Eggen ¹, Kenneth Caidahl ^2,³, Lars Dähne ⁴, Rune Hansen ¹

Editor: Joseph Donlan⁵

PMCID: PMC9632906 PMID: 36327225

Abstract

Contrast enhanced ultrasound is a powerful diagnostic tool and ultrasound contrast media are based on microbubbles (MBs). The use of MBs in drug delivery applications and molecular imaging is a relatively new field of research which has gained significant interest during the last decade. MBs available for clinical use are fragile with short circulation half-lives due to the use of a thin encapsulating shell for stabilization of the gas core. Thick-shelled MBs can have improved circulation half-lives, incorporate larger amounts of drugs for enhanced drug delivery or facilitate targeting for use in molecular ultrasound imaging. However, methods for robust imaging of thick-shelled MBs are currently not available. We propose a simple multi-pulse imaging technique which is able to visualize thick-shelled polymeric MBs with a superior contrast-to-tissue ratio (CTR) compared to commercially available harmonic techniques. The method is implemented on a high-end ultrasound scanner and in-vitro imaging in a tissue mimicking flow phantom results in a CTR of up to 23 dB. A proof-of-concept study of molecular ultrasound imaging in a soft tissue inflammation model in rabbit is then presented where the new imaging technique showed an enhanced accumulation of targeted MBs in the inflamed tissue region compared to non-targeted MBs and a mean CTR of 13.3 dB for stationary MBs. The presence of fluorescently labelled MBs was verified by confocal microscopy imaging of tissue sections post-mortem.

Introduction

Contrast enhanced ultrasound imaging is an important diagnostic tool in clinical practice, where microbubbles (MBs) are injected intravenously to enhance the signal from the blood pool. The MBs in clinical use consist of thin lipid or protein shells encapsulating the gas core, and applications for imaging such MBs are based on harmonic methods, taking advantage of the highly flexible shell which facilitates resonant and nonlinear backscattering at low incident mechanical indices [1–5]. Pulse inversion (PI), amplitude modulation (AM) or combinations of the two are most common harmonic methods on clinical ultrasound scanners. In the later years substantial research effort has been focused on functionalizing MBs, either by incorporating drugs for enhanced drug delivery [6–8] or attaching targeting ligands for active molecular targeting towards disease-specific biomarkers [9–11]. The shell properties of functionalized MBs may become very different from the thin-shelled MBs, and hence new imaging methods optimized for thick-shelled MBs may be required. A thick shell (>100nm) will in general have higher viscosity and therefore introduce higher damping of the bubble oscillation compared to a thin shell. Compared to ultrasound imaging of thin-shelled MBs, higher mechanical indices are typically required to drive thick-shelled MBs into nonlinear oscillations [12]. Contrast harmonic imaging methods typically rely on transmit pulses with very low mechanical indices, sufficient to invoke nonlinear scattering from thin-shelled MBs while suppressing harmonic components, as well as the strong fundamental component, from soft tissue. Tissue harmonic imaging methods, with higher mechanical indices, have been in clinical use for more than two decades [13] and the contrast-to-tissue-ratio (CTR) will typically be destroyed when increasing the mechanical index for imaging of MBs. We propose a novel multi-pulse imaging method where imaging pulses are combined with intermediate manipulation pulses at high mechanical index. Radiation force both from the imaging and manipulation pulses, in addition to possible nonlinear effects from changes in the shell result in contrast enhancement of thick-shelled MBs with adequate tissue suppression. Contrast enhancement methods involving the use of radiation force have been proposed previously, especially focusing on pushing targeted MBs towards the vessel wall to enhance the targeting effect [14–16]. The radiation force imposed by the proposed method might also enhance the targeting, but it is also the primary source of the MB detection signal.

A thicker and more stable shell can increase the in vivo stability of the MBs and increase the circulation half-life [17, 18] which currently is limited to 1–3 minutes with thin-shelled MBs [19, 20]. An increased MB circulation time is especially important for drug delivery and molecular imaging applications. For drug delivery, it will contribute to enhanced accumulation of drug in target regions, when the MBs themselves are loaded with drugs, and to prolonged duration of MB oscillations within a target region with co-injection of drugs and MBs. For molecular imaging, it will contribute to enhanced exposure of target tissues to the targeted MBs.

Molecular imaging utilizes the principle of labelling an image sensitive vector with a ligand that can bind specifically to receptors expressed by cells in the body, and the purpose is to use a non-invasive method to discover alterations in the physiology at a molecular level. Ultrasound molecular imaging can primarily be performed using targeting towards biomarkers expressed by structures within the lumen of blood vessels, such as the luminal surface of the endothelial cells, and several research groups have successfully demonstrated an increased accumulation of targeted MBs in preclinical experiments in tumor tissue [21–23], inflammation [24] and in atherosclerosis [25, 26]. BR55 (Bracco Research inc.) with specificity towards VEGFR2 is the first targeted ultrasound contrast agent that have reached clinical trials. Breast, ovarian and prostate cancer were the inclusion criteria in the first trials [27, 28], and there is ongoing recruitment on pancreatic tumors (NCT03486327 at www.clinicaltrials.gov).

As a proof-of-concept study of the novel MB imaging technique developed for imaging thick-shelled MBs, and to show their suitability for ultrasound molecular targeting, a simple and fast inflammation model suited for soft tissue is proposed. Inflammation is a basic, yet complex dynamic response of the vascularized body systems to any harmful stimuli. The classical macroscopic signs of inflammation are pain, heat, redness, swelling and loss of function. However, when studying inflammation on the microscopic level, several cascades of cellular and microvascular reactions are present, which can be exploited in targeted drug delivery and molecular imaging [29]. Induction of a sterile inflammation can be achieved by injecting zymosan in the organ of interest. Zymosan is a substance derived from the cell wall of the yeast Saccharomyces cerevisiae, and has previously been used to study peritonitis [30, 31], arthritis [32], multiple organ dysfunction syndrome (MODS) [33], lung disease [34] and the effect of anti-inflammatory drugs [35]. The induction of a local inflammation by zymosan injection in muscle tissue has previously been used to investigate gamma camera imaging of inflammation [36], and a similar approach for induction of inflammation in the hind leg using lipopolysaccharide (LPS) has been reported [24].

In this paper, we present a proof-of-concept study where a novel ultrasound contrast imaging technique, which is suitable for imaging thick-shelled MBs is demonstrated. The new method is tested in vitro and in vivo, and we provide imaging results from a zymosan-induced inflammation model in muscle tissue in rabbit and compare the amount of targeted and non-targeted MBs in inflamed and healthy tissue.

Materials and methods

Microbubble preparation

The synthesis of the thick-shelled polymeric MBs has been described by Cavalieri et al [37, 38]. The MBs were manufactured at Surflay Nanotec GmbH, Berlin, Germany, and they have a shell made of polyvinyl alcohol (PVA) [39, 40]. These MBs are obtained by foaming a solution of PVA previously oxidized with sodium metaperiodate. The PVA chains are cross-linked during reaction occurring at the water/air interface. Resulting MBs have an average diameter of 3 μm with a shell thickness of about 200 nm and an approximate concentration of 10⁹ MB/ml. Concentration is measured by counting MBs in a Neubauer Cell under a confocal laser scanning microscope. A microscopy image of the MBs and their size distribution is shown in Fig 1.

In the molecular ultrasound imaging experiments MBs were functionalized by letting remaining aldehyde groups in the PVA shell react with Aminoguanidine hydrochloride yielding a positive charge. This enables a further layer-by-layer-coating [41, 42] by (Poly(styrenesulfonate/Poly(ethyleneimine)₂/Polystyrenesulfonate. The second Polyethylenimine layer was labeled with the fluorescent dye Tetramethylrhodamine-isothiocyanate (TRITC). These MBs with only TRITC in the shell are called tritc-MBs. For the second type of the MBs a further layer of biotin labelled Poly(allylamine) was coated to the surface. On the biotinylated surface streptavidin was bound, enabling attachment of biotinylated polyclonal rabbit anti-ICAM-1 (CD54) (bs-4615R-Biotin), purchased from Bioss Inc., Woburn, MA, USA. Here 60 μg of biotinylated antibody was added to 1.5 ml of MBs at a concentration of 10⁹ MBs/ml. Before injection into animals the MB solution was diluted in saline to 2 x 10⁸ MB/ml. The MBs with both TRITC and streptavidin are called strep-tritc-MBs.

In-vitro ultrasound imaging

Due to their thick shell, the PVA MBs used in the current experiments were not well imaged by conventional harmonic imaging schemes based on PI or AM. An experimental multi-pulse technique (as described below) was therefore implemented, and imaging results from PI and AM were compared with the new technique. The novel multi-pulse technique was implemented on a GE Vivid E9 scanner (GE Vingmed, Horten, Norway) modified for research, and the 11L linear transducer was used. Comparison to clinically available techniques was also performed, and a GE Vivid E95 in clinical mode with the 9L transducer was used for PI and AM imaging.

In the multi-pulse scheme, several pulses were transmitted in each beam direction. The first and the last pulse were identical, and were used for imaging, whereas one or several intermediate pulses could be included for additional manipulation of the MBs, as illustrated in Fig 2. The intermediate pulses could have a different center frequency and number of oscillations, however, in the presented experiments the transmitted intermediate pulses were equal to the imaging pulses due to limited access to the transmit setup files of the scanner. Contrast enhanced images were formed by subtracting the echoes resulting from the imaging pulses (first and last) transmitted along each image line.

Fig 2 — For each scanline, two imaging pulses and several manipulation pulses were transmitted. Upon reception the last pulse was subtracted from the first, resulting in a contrast enhanced signal for that particular scanline. The pulse length, frequency and amplitude of the intermediate manipulation pulses could be different from the imaging pulses, but in this work, all transmitted pulses along one scanline were equal due to practical reasons.

Transmit frequencies at 8 and 9.5 MHz were tested, and acquisition schemes with MI of 0.7 and 1.0, and pulse lengths of 3 and 5 half periods were compared. Images were recorded and stored in the scanner archive, and analysis on recorded images were done in the GE software EchoPAC (GE Vingmed, Horten, Norway). In addition, RF-data were recorded and processed in an in-house Matlab-environment.

Ultrasound imaging tests, method development and optimization with the thick-shelled polymeric MBs were performed with an in-vitro setup consisting of a tissue mimicking flow phantom from ATS, model 524 (ATS Laboratories, Norfolk, VA, USA) and a peristaltic pump. The ultrasound transducer was fixed with a clamp and MBs were imaged both during low-speed flow and when the pump had stopped and the MBs were stationary. All analysis of CTR was done on images recorded after the pump was stopped.

In-vivo experiments

Chemicals

Zymosan A (Sigma Aldrich Co., St. Louis, MO, product no: Z4250) was suspended at 1% in 0.15 M sodium chloride and placed in a boiling water bath for one hour, followed by centrifugation for 30 minutes at 4000 rpm. The supernatant was discarded, and the residue suspended evenly in 0.9% NaCl to a concentration of 35 mg/ml.

Animal preparation and experimental procedures

Experimental procedures with female New Zealand White rabbits were conducted in compliance with protocols approved by the Norwegian National Animal Research Authorities (Protocol number: FOTS-5360) and all animals were acclimatized for at least one week before the experiments started. Illustrations of the timelines and experimental procedures are shown in Fig 3.

Fig 3 — An intramuscular injection of zymosan was given at t = 0 (indicated by a green syringe), and MBs were administered i.v. at t = 2 hours to the first rabbit and at t = 24 hours for the second and third rabbit (indicated by a pick syringe). Ultrasound imaging was performed on all three days and all rabbits were euthanized 72 hours after the zymosan injection.

Rabbits (Hsdlf:NZW) were purchased at 15 weeks of age from Harlan Laboratories (The Netherlands). During acclimatization, the rabbits were housed as group of three in a rabbit cage system supplied by Scanbur, type EC3, and during the experimental period they were housed individually. Commercial diet and free access to hay and water were provided. They were purchased with status as specific pathogen free (SPF), but were housed in a non-SPF section since the Comparative Medicine Core Facility at NTNU does not offer SPF conditions for rabbits. The rabbits were sedated with a subcutaneous injection with a solution of fentanyl 0.2 mg/ml and fluanisone 10 mg/ml (Hypnorm^®, VetaPharma Ltd, Leeds, UK) at a dosage of 0.3 ml/kg and were given an intravenous injection of midazolam (5 mg/ml (Midazolam B. Braun, Melsungen, Germany) at a dosage of 2 mg/kg) through a venflon in the lateral ear vein to obtain full anesthesia. The anesthesia was maintained during the experiment with inhalation of 2% isoflurane gas in oxygen to keep the anesthesia at a surgical level. The venflon was kept in place for the administration of contrast agents. On the last day of experiments, the rabbits were euthanized by an intravenous administration of 5 ml of Pentobarbital 100 mg/ml.

Three rabbits were included in the experiment, and sedation and anesthesia were repeated on each day for four days. The fur on both hind legs was removed using an electrical clipper. On the first day, a sterile inflammation was induced by an ultrasound-guided intramuscular injection of 1 ml zymosan, at a concentration of 35 mg/ml in the area of the biceps femoris muscle in the left hind leg. The right hind leg was the negative control. Two rabbits were given intravenous injections of approximately 5x10⁸ strep-tritc-MBs with anti-ICAM1 following the zymosan injection. The first rabbit was given the MBs 2h after zymosan injection, whereas the second rabbit was given the MBs 24 h after zymosan injection. The third rabbit was given an injection of approximately 5x10⁸ tritc-MBs 24 h after the zymosan injection. The development of the inflammation and the MBs were imaged by ultrasound at 24h, 48h and 72h. All rabbits were sacrificed at 72h. Post-mortem, tissue biopsies were excised from both the left and right hind leg of all three rabbits. The samples were embedded in O.C.T. (Tissue-tek, Sakura Finetek, The Netherlands) snap frozen and stored at -80 degrees until sectioning.

A pain scoring sheet was logged daily for all the animals to assess their general condition and clinical symptoms. To ensure minimal pain the rabbits were given daily subcutaneous injections of 0.05 mg/kg and 0.03 mg/kg of buprenorphine (Temgesic^®; Reckitt Benckiser, UK) respectively.

Molecular ultrasound imaging

B-mode and contrast-enhanced ultrasound images of the hind legs of the rabbits were acquired using a GE Vivid E9 scanner modified for research and an 11L linear transducer. The simple multi-pulse imaging technique was used to enhance the MBs. Images were acquired during MB injection and at 24h, 48h and 72h after the inflammation was induced. A frequency of 9.5 MHz and MI of 1 was used in the contrast images, and 12 MHz and MI of 1 was used to acquire the B-mode images. Preliminary imaging showed that the inclusion of intermediate pulses caused increased tissue signal in form of flashing. In order to minimize the flashing from tissue and increase the framerate, the in vivo recordings were done without any intermediate pulses, hence only the two imaging pulses were transmitted along each beam line. The CTR was calculated based on traces exported from the Q-analysis tool in the EchoPAC software. In each recording which was analyzed, 8 circular regions of interest (ROIs) with diameter of 0.5 mm were drawn in positions where MBs were present in some of the recorded frames, and the CTR was calculated as the difference between the peaks in signal level found when an MB was present, and the background level in the same position, when no MB was seen. All datapoints below -50 dB were considered to come from tissue and all peaks above -43 dB were considered to be caused by MBs. Multiple MBs were typically detected in each of the ROIs. Ultrasound video files used for CTR analysis and signal level traces from all ROIs are provided in the data repository.

Histology and microscopy imaging

Sections with a thickness of 4 μm and 25 μm were made from the rabbit muscle tissue blocks. The 4 μm sections were HES-stained and the 25 μm sections were mounted with Vectashield mounting medium with DAPI (Vector Laboratories Inc., Burlingame, CA, USA). A selection of the tissue sections was examined by microscopy techniques.

HES stained tissue sections were examined with a Nikon white light microscope with objectives from 4x magnification. Tissue sections of 25 μm with DAPI stained cell cores were examined with a confocal laser scanning microscope (CLSM) (Zeiss LSM510, Germany), where a Helium-Neon (HeNe) laser line at 543 nm was used to detect the fluorescent dye TRITC in the MBs, with emission in the range 560 nm to 615 nm, and a multi-photon Titanium-Sapphire (TiSp) laser line at 750 nm was used to detect the DAPI bound to DNA in the cell nuclei, with emission in the range 420 nm to 470 nm. A 20x /0.5 air objective was used in combination with the tile scan function to obtain images of the whole tissue sample. Each single image had a resolution of 512x512 pixels and covered 450x450 μm. A total of 11 section from inflamed tissue and 7 sections from healthy tissue from the three rabbits were examined. MBs were automatically counted in ROIs of 1 x 1 mm² with a custom-made Matlab script (R2019a, The MathWorks Inc., Natick, MA, USA), and the number of MBs in the ROI with the maximum amount of MBs in each section were registered and stored.

Results

In vitro ultrasound imaging

Ultrasound image optimization performed using a tissue mimicking flow phantom showed that with the traditional PI and AM techniques, which are implemented on most commercial scanners, it was not possible to both visualize the thick-shelled MBs properly and suppress the signal from the tissue. This is shown in the panel A and B of Fig 4. With PI it was only at an MI of 0.1 that the tissue signal was adequately low, whereas for AM the tissue suppression was satisfactory also at MI of 0.2. However, the MBs were not possible to detect at either of these pressure levels. An increase of the MI to 0.6 was needed to achieve adequate MB detection, but the tissue signal at such high MI also became high. AM had better tissue suppression in shallow parts of the image, but in the deeper parts, the signal intensity from the tissue mimicking material was in the same range as the MB signal. Hence, neither of the methods were suitable for imaging the thick-shelled MBs in small blood vessels and capillaries. The dynamic range was the same for all PI images, but the gain was adjusted for each MI setting to get a similar level of electronic noise within the flow tube at all three pressure levels.

Images acquired with the multi-pulse detection method proposed in this work are shown in the two lower panels of Fig 4. In a stationary phantom, the subtraction of the second imaging pulse from the first, resulted in images where the tissue signal was completely suppressed. When only the two imaging pulses were used, the signal from the MBs was clearly seen inside the flow channel, but including one or two intermediate manipulation pulses between the imaging pulses, enhanced the MB signal even more. In the images which included manipulation pulses the improved detection can be seen by a somewhat higher intensity and an increased size of the MBs. The tissue signal is not affected by the manipulation pulses in this setup.

When recording RF data from the scanner, an analysis of the received pulses along each scanline is possible. An example of RF data analysis from a recording at 9.5 MHz with an MI of 1.0 and pulse lengths of 3 half periods is shown in Fig 5. The backscattered signal from many of the MBs is much lower than the strongest tissue signal, as can be seen in the B-mode image in Fig 5A, which is based on the first imaging pulse only. Subtracting the last pulse from the first results in a contrast enhanced image which is shown in Fig 5B. When comparing the amount of MBs in the image in Fig 5A and 5B, it is apparent that not all MBs become visible with this method. However, the MBs that are detected have a high CTR. In Fig 5C–5F, the MB marked with a white arrow in Fig 5B is further analyzed. The received imaging pulse along the scanline through the MB is displayed in Fig 5C, and a zoom-in on the backscattered signal from the MB is shown in Fig 5D. Subtracting the last imaging pulse from the first, results in a subtraction signal where the stationary areas of the scanline are close to zero and the signal from the MB is enhanced due to a small translation as shown in Fig 5E (and zoomed in in Fig 5F). Analysis of the signal from this particular MB gives a detected delay between first and last pulse of 5 ns, corresponding to a translation of 3.9 μm and a resulting CTR of 22.6 dB.

Fig 5 — The analysis is based on RF data from in-vitro imaging with a GE Vivid E9 scanner. A: B-mode image of stationary MBs in a flow channel through a tissue mimicking phantom. The image is based on only the first imaging pulse. B: Contrast enhanced image resulting from subtracting the last imaging pulse from the first. The image line which goes through the center of the MB marked by the white arrow is analyzed further. C: Received signal from imaging pulses along the indicated image line. D: Received signal from the MB marked by the white arrow (zoomed in on the signal marked by a black box in panel C). The second pulse is delayed by approximately 5 ns compared to the first pulse. E: The resulting signal when the second imaging pulse is subtracted from the first. F: Resulting MB signal after subtraction (zoomed in on the data marked by a black box in panel E).

A systematic collection of RF data from images recorded with two intermediate pulses at 8 and 9.5 MHz, with MI of 0.7 and 1 and pulse lengths of 3 and 5 half periods was analyzed. When subtracting the second imaging pulse from the first, the delay caused by a translation of the MBs of 2–5 μm resulted in a CTR of up to 23.1 dB. The backscattered signals from the 6 to 10 brightest MBs in each image were analyzed, and the detected mean delay between the first and last imaging pulse along the scanline through the MB and the mean CTR are presented in Table 1. In addition to the results in Table 1, RF data from images recorded at 8 MHz without manipulation pulses was also recorded. At an MI of 1 the mean delay when using 3 hp and 5 hp imaging pulses was 2.2 ± 0.7 ns and 2.3 ± 0.8 ns respectively. The corresponding mean CTR was 13.0 ± 1.9 dB and 21.7 ± 2.1 dB for 3 and 5 hp. At MI = 0.7 there were hardly any visible MBs in the contrast image.

Table 1. Mean delay and CTR ratio in images recorded at 8 and 9.5 MHz.

Two intermediate manipulation pulses between the imaging pulses were used in the experiments.

Frequency	8MHz				9.5MHz
Pulse length	3 hp	5 hp	3 hp	5 hp	3 hp	5 hp	3 hp	5 hp
MI	1.0		0.7		1.0		0.7
Mean delay ± std [ns]	5.0 ± 0.9	5.3 ± 0.6	2.6 ± 0.4	2.4 ± 0.5	4.4 ± 0.7	4.2 ± 0.5	2.2 ± 0.4	2.4 ± 0.3
Mean CTR ± std [dB]	23.1 ± 2.0	20.4 ± 3.2	16.0 ±3.2	13.9 ± 3.3	22.4 ± 2.5	19.2 ± 1.7	12.9 ± 1.9	11.7 ± 1.6

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Abbreviations: contrast-to-tissue ratio, CTR; half period, hp

The high MI setting (MI = 1.0) was the most important contributor to an increased CTR, with a 7–10 dB increase compared to recordings at MI of 0.7 and otherwise equal settings. Pulse length and frequency are of less importance. In order to get the best image resolution, the highest frequency and shorter pulse length were chosen as initial settings for the in vivo experiments, i.e. 9.5 MHz and 3 half periods. However, due to motion, and flashing artefacts when using the intermediate manipulation pulses in vivo, the in vivo results were recorded without manipulation pulses at MI = 1.

In-vivo ultrasound imaging

B-mode ultrasound images of healthy thigh muscles generally showed a quite uniform signal intensity from the muscle tissue. In the thigh muscle with induced inflammation, a hypoechoic (dark) region was evident 24 hours post-injection of zymosan. The extent of architectural changes was different between animals, but the hypoechoic areas detected in one leg was recognizable the following days and a general observation was that these areas became more extensive with time. As shown in Fig 6 the hypoechoic (dark) area was clearly detectable after 24 hours and increased in size after 48 hours and 72 hours. Upon injection of MBs, the circulating MBs were detected by the novel multi-pulse contrast enhanced ultrasound imaging method in both small and large vessels within the thigh muscle for several minutes. CTR analysis of the recording from each of the rabbits just after the MB injections resulted in a tissue signal level of -53.1 ± 2.0 dB and contrast signal of -36.3 ± 4.6 dB, giving a CTR of 16.8 dB for flowing MBs. Ultrasound videos of inflow of MBs are available in the data repository. When imaged 24h, 48h and 72h after the injection, no circulating MBs were seen; however, a large number of strep-tritc-MBs with anti-ICAM-1 were detected along the perimeter of the inflamed area (Fig 7D). By contrast, the tritc-MBs without active targeting were not found close to the inflammation, but some tritc-MBs were detected within the areas of healthy muscle tissue (Fig 7F). No apparent MB destruction was observed by ultrasound imaging, even though the MI during imaging was 1.0. The signal from the non-circulating MBs was stable over the entire imaging period, and CTR analysis of one recording at each timepoint for each of the rabbits resulted in the same signal level from tissue as before and an average contrast signal of -39.8 ± 2.5 dB, giving a CTR of 13.3 dB for static MBs. A total of 172 local signal peaks representing MBs were found in the ROIs from images recorded just after MB injection and 205 local peaks were found in the ROIs in images 24h, 48h, and 72h after MB injection. Figures showing examples of the traces representing signal within selected ROIs and the distribution of CTR levels from all the detected MBs from the various recordings can be found as S1 File.

Fig 6 — Image a) shows healthy muscle tissue, and the white arrow points to the tip of the needle injecting the zymosan. Images b), c) and d) shows the hypoechoic region, which developed over the three days following the zymosan injection. The imaging depth is 3 cm and dynamic range is 69 dB.

Fig 7 — Images A and B are from Rabbit 1, which was injected with strep-tritc-MBs with anti-ICAM1 2 hours after the onset of the inflammation. The images shown are recorded 22 hours after MB injection, and the area of inflammation is marked with white arrows. Images B and C are from Rabbit 2, which was injected with strep-tritc-MB with anti-ICAM1 24 hours after the inflammation was induced, and the images are recorded 48 hours after MB injection. The inflammation area and accumulation of MBs at its perimeter are marked by white arrows in the images. Some MBs are also detected further from the inflammation (right side of (D). In images E and F examples from Rabbit 3 are shown. The animal was given tritc-MB (MBs without streptavidin and antibodies) 24 hours after the inflammation was induced, and the images are recorded 48 hours after MB injection. The contrast enhanced image shows some MBs in the muscle tissue at the far-right end (red arrow), but no accumulation of MBs around the inflammation area (white arrow). B-mode images are recorded at 12 MHz and shown with a dynamic range of 69 dB and contrast enhanced images are recorded with imaging pulses of 9.5 MHz and a dynamic range of 36 dB.

Muscle tissue necropsy

Upon post-mortem dissection, an accumulation of a pale, yellow viscous exudate was found in the inflammatory region. The tissue in the parts of the muscle surrounding the inflammation was darker in color, hence indicating that the blood flow was enhanced as a sign of an inflammatory reaction.

Ex vivo analysis

In HES-stained tissue sections, evident changes could be seen in the muscle tissue from the hind leg where the inflammation had been induced, in all animals. The accumulation of inflammatory cells was extensive and infiltrated areas between the surrounding muscle fibers, as seen in Fig 8A. The exudate found in the tissue with inflammation consists of a protein-rich fluid, dead leukocytes and cellular debris. The inflammatory cells consisted of mainly granulocytes, but also mononuclear cells were present. In areas of the tissue sections located 3–5 mm away from the accumulation of inflammatory cells, the muscle tissue morphology appeared to be normal.

Fig 8 — Confocal microscopy images (B, C and D) are from a 25 μm thick tissue section. The cell nuclei are stained with DAPI and shown in green and the MBs containing TRITC are shown in red. Two selected regions (0.5x0.5mm) from the confocal image (B) are displayed in detail in C and D. The white boxes represent areas of pus-filled region displayed in C, and the green boxes mark an area approximately one millimeter away from the pus-filled region and contains muscle tissue with only slight infiltration of inflammatory cells. Boxes with full lines represent the area in the confocal image and corresponding positions in the HES-stained section are marked with dashed boxes.

The confocal microscopy examination of tissue sections from the inflamed and healthy muscle tissue from rabbits that were given strep-tritc-MBs with anti-ICAM-1 and tritc-MB without targeting confirmed the findings from the contrast enhanced ultrasound images. A large number of MBs were found in the areas close to the pus-filled regions compared to the areas further away. In the inflamed muscle tissue of the rabbit that was given tritc-MBs without any active targeting and in the healthy muscle tissue, hardly any MBs were detected by confocal microscopy imaging. An example image from the rabbit that was given targeted strep-tritc-MBs 2h after induction of inflammation is shown in Fig 8. Enhancement of an area within the active inflammation (Fig 8C) and an area 1 mm away from the inflammation (Fig 8D) show that MBs are mainly found in close vicinity to the pus-filled region. The MBs can be seen as red dots in Fig 8C.

MBs were counted in a total of 11 tissue sections from inflamed tissue and 7 sections from healthy tissue, and the results are shown in Fig 9. The highest number of MBs was found in the inflamed tissue from the rabbit which was given the strep-tritc-MBs with anti-ICAM1 only two hours after the inflammation was induced. In the healthy tissue the mean number of MBs/mm² varied from 0 to 1.3, and no apparent difference was found between targeted and non-targeted MBs.

Animal health

The animals did not display any visible signs of pain or distress, and the body weight did not change significantly during the experimental period. The inflammation was not detectable by visual inspection only, but through palpation, areas that were firmer than a healthy thigh muscle were found in the areas where the zymosan had been injected.

Discussion

In this study, we have demonstrated a simple multi-pulse technique for imaging thick-shelled MBs both in vitro and in vivo in an inflammation model for soft tissue, where its suitability for use in molecular ultrasound imaging also is shown.

Ultrasound imaging technique

When comparing the proposed multi-pulse technique with the commercially available PI and AM techniques, we showed that the multi-pulse technique was superior in visualizing thick-shelled MBs. PI utilizes the fact that scattering from commercially available thin-shelled MBs, like SonoVue, will be nonlinear even at very low mechanical indices (below 0.1) when the driving frequency is below or in the vicinity of the MBs resonance frequency. For MBs with a thick shell or a less flexible shell, and when the driving frequency is above the resonance frequency, higher mechanical indices are required to obtain adequate nonlinear scattering from the MBs. Higher mechanical indices will result in nonlinear wave propagation through the tissue causing harmonic components to be backscattered from tissue. With harmonic detection techniques, the differentiation between MBs and tissue will then be compromised, as was shown in the upper panels of Fig 4. AM could be expected to perform better than PI when imaging thick-shelled MBs, since it relies on the difference in MB response between two pulses with different amplitude [43]. When imaging thin shelled MBs this difference is caused by MB oscillation. In a situation with thick-shelled MBs one could envision that a high amplitude pulse could cause a backscatter signal, e.g. due to shell buckling [44], whereas the low amplitude pulse did not, and by that detecting the presence of MBs. However, the in vitro imaging showed that both PI and AM only could detect MBs at MI levels which were too high for the tissue suppression to work properly, whereas there was no detectable MB signal at low MIs.

When using the multi-pulse technique, the MB-signals arise due to a difference in the received signals between the two imaging pulses transmitted along each scan line. The most important contribution to the detection signal appears to be a delay between the two resulting backscattered signals due to a small physical translation of the MBs caused by the radiation force from the incoming pulses. In the example shown, the detected delay was in the range 2–5 ns, which corresponds to a translation of a few micrometers. In the in vitro setup, the tissue mimicking material is stationary and the MBs are in a tube with a diameter much larger than the MB diameter. In this idealized situation a CTR of up to 23 dB was achieved. In an in vivo situation, the tissue is not expected to be stationary, and the MBs in small vessels or capillaries will have limited ability to undergo a translational motion. Hence the CTR was expected to be lower. In the example of MBs imaged in the vicinity of the inflammation area a mean CTR of 13.3 dB was obtained.

Including several manipulation pulses or a manipulation pulse of longer duration between the imaging pulses may contribute to detection of a larger fraction of MBs and stronger signal from the MBs as long as the MBs are able to physically translate, but prolonged time between the two imaging pulses will also result in increased tissue signal unless the tissue is stationary. Tissue motion may especially be a problem in the abdominal and thoracic areas, where breathing and the beating heart may cause strong artefacts if no motion compensation is included in the implementation. In the current work, a relatively stationary organ was imaged (the thigh). Artefacts due to motion were therefore minimized and the use of motion compensating algorithms was not required, however, there is still some flashing effects which may be caused by probe translation and can be seen in the videos in the data repository.

Another potential contribution to the detection signal is due to a change in amplitude or pulse form between the two resulting signals scattered from the MBs. Changes in amplitude or pulse form may be caused by a partial rupture, shell perturbation or deflation of the MB due to the oscillations caused by the incident pressure pulses. We have not been able to verify the presence of such effects in the current work, but the same kind of MBs have been studied with high speed optical imaging during ultrasound exposure, and it was found that at low frequencies (2–4 MHz) the shell buckled and that the gas was pumped out of the MBs when bursts of 10 cycles were applied [45]. However, several consecutive pulses at high MI were needed to effectively rupture the MBs, and the pulses used in the presented experiments are only 3–5 half cycles, and at a much higher frequency (9.5 MHz), so the effect on the MBs might be somewhat different than what was observed in the experiments of Kothapalli et al. A full destruction or deflation of the MB during the insonation of short imaging pulses is unlikely, and also backed by the observations during the imaging.

The multi-pulse technique proposed in this study is implemented with a transmit frequency at 8 and 9.5 MHz but can in principle be used at any frequency range with suitable transducers and ultrasound scanners. A modified detection sequence based on power Doppler using short (e.g., 1-cycle) pulses of suitable ensemble length, for optimal spatial resolution and clutter/tissue filtering, with intermediate manipulations pulses of suitable frequency and duration, for inducing bubble decorrelation, is probably the optimal implementation of this method. This will be pursued in future projects.

The inflammatory reaction

The accumulation of pus in the region where the zymosan was injected is a sign that leukocytes have been recruited from the circulation and transmigrated into the tissue where the initial damage occurred. This recruitment is a result of production of cytokines and subsequent inflammatory specific changes in the endothelium that allows for circulating immune cells to enter the injured area and, if possible, eliminate the cause and repair the damage. The ultrasound images show that the onset of the inflammatory reaction was rapid, and that the amount of pus and extent of infiltration of inflammatory cells in the muscle tissue increased with time. It is a great advantage, both for animal health and cost of experiments, that the time from induction of disease to the point where the molecular imaging can be performed is short.

Molecular ultrasound imaging

The polymer MBs used in these experiments have longer blood circulation time compared to commercially available lipid MBs [46], and the ultrasound images of MBs in the inflamed area show that single MBs can be detected even days after injection. Active targeting and hence attachment of MBs to endothelial cells might hinder or diminish the translation effect due to radiation force and it will probably also reduce the fraction of MBs being detected. However, in most situation, a component of the radiation force will act in the flow direction of the capillary, which could cause the MBs to detach from the endothelial cells. During imaging we observed that the detected MBs could disappear after being imaged for some time (typically ranging from about 5 seconds to several minutes) potentially indicating some sort of rolling and eventual detachment.

Studies of similar PVA MBs containing superparamagnetic iron oxide (SPIONs) in the shell as MR imaging contrast agent, showed that the MBs were phagocytized by macrophages, mainly in the liver and spleen, and that the process took weeks [17, 18]. This is also supported by the work of Härmark et al [46], where PVA MBs were found only inside vasculature, but often taken up by or in the vicinity of macrophages. Successful targeting of ICAM-1, VCAM- 1, and e-selectin using layer-by-layer MBs has been shown in human and murine endothelial cells and in a zymosan induced peritonitis model [42].

Results from the histological examination show that the rabbit given strep-tritc-MBs with anti-ICAM-1 only 2 hours after induction of the inflammation had a higher concentration of MBs in the vicinity of the pus-filled regions compared to the rabbit given the injection 24 hours after inflammation induction. This indicates that the total circulation time (70 hours vs 48 hours) resulted in a higher number of MBs attaching to biomarkers. Another possible reason is that more targeted biomarkers were available 2 hours post induction than at 24 hours after the inflammation was induced. However, with the limited number of animals included in the proof-of-concept study it is not possible to draw a conclusion in this matter. The tritc-MBs without anti-ICAM-1 were not found in the vicinity of the inflamed region, indicating that the active targeting with anti-ICAM-1 was the main cause for the high number of MBs close to the inflammation. However, since the tritc-MBs did not contain streptavidin in the shell, the possibility of an unspecific binding of the streptavidin in the strep-tritc-MBs to the inflamed regions cannot be excluded. The strep-tritc-MBs detected by ultrasound in healthy tissue 24h and 48h after injection may also indicate a level of a general unspecific binding due to the streptavidin on the MB shell. Some tritc-MBs were also detected within healthy tissue in the contrast enhanced ultrasound images. However, based on confocal microscopy evaluation of tissue section the number of MBs within normal tissue was very low for all three rabbits. The increased blood flow in the inflamed area may also contribute to the high number of MBs found, compared to normal muscle tissue.

Conclusion

A simple multi-pulse imaging technique for contrast enhancement of thick-shelled MBs has been described and implemented on a high-end ultrasound scanner. Comparison to traditional PI and AM techniques showed superior in vitro detection quality. In vivo, the technique was demonstrated in molecular ultrasound imaging of a soft tissue inflammation model in rabbit. MBs targeted towards the inflammatory marker ICAM-1 were injected and visualized in the muscle tissue upon injection and up to 70 hours after injection. An increased number of MBs in the perimeter of the inflamed area was demonstrated both on contrast enhanced ultrasound images and on histological examinations of excised tissue.

The presented soft tissue inflammation model enables molecular imaging of a relevant inflammatory biomarker within a day after the induction of inflammation and does not lead to apparent pain or discomfort for the animals. It is a flexible model with respect to both anatomy and species, which makes it highly relevant for testing and optimization of new contrast agents and contrast agent imaging techniques.

Supporting information

S1 File

(DOCX)

Click here for additional data file.^{(216.8KB, docx)}

Acknowledgments

The procedures involving laboratory animals were performed at the Comparative Medicine Core Facility (CoMed), Norwegian University of Science and Technology (NTNU), Trondheim, Norway. The embedding of tissue samples, preparation of tissue sections, HES staining was performed at the Cellular & Molecular Imaging Core Facility (CMIC), NTNU. Thanks to dr. Per Stenstad at SINTEF Industry, Trondheim, Norway for attaching antibodies to the MBs, and to GE Healthcare for allowing us to implement an experimental contrast imaging scheme on the GE Vivid E9 system.

Data Availability

Underlying data is fully available at the following repository: https://doi.org/10.11582/2021.00072.

Funding Statement

SB: Liaison Committee between the Central Norway Regional Health Authority (RHA) and the Norwegian University of Science and Technology (NTNU) SB and RH: Research Council of Norway (240410/F20) SB, SE, LD, KC and RH: European commission (7th framework program, 3MiCRON (245572) project).

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PLoS One. doi: 10.1371/journal.pone.0276292.r001

Decision Letter 0

Guy Cloutier

2 Feb 2022

PONE-D-21-28949A multi-pulse ultrasound technique for imaging of thick-shelled microbubbles demonstrated in vitro and in vivoPLOS ONE

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Three expert reviewers provided detailed and relevant comments. Among many other issues, some "solvable" flaws were identified. It is clear that the framework of the comparative evaluation with different sequences is seriously questioned since the acoustic pressure used and frequencies are different. Additional experiments would be required to better support your statements and quantitative analysis would also be required in vivo. Dr Cloutier

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Reviewer #1: The authors present a succinct piece of work using a new imaging sequence that is tailored for stiff microbubbles. The new imaging sequence is very similar to amplitude modulation, in that it is measuring the difference between the response of a bubble population over time, which can be due to translation, bubble rupture, or other effects. I would have liked to see a direct comparison between the new sequence and AM, not PI as the authors did. The comparison between PI, indeed, was taken at entirely different frequencies and pressure amplitudes, and I’m not sure it’s a fair comparison. The work seemed to be conducted robustly, and in that regard the paper deserves eventual publication. I would perhaps suggest to the authors that, in addition to my comments below, the authors comment on the similarly between AM, PI and their new sequence since the physics isn’t different. This approach essentially is trying to get away from bubble vibration at all (since they don’t vibrate very well at low MI) and is trying to focus on signal ‘difference’ due to motion, etc. What is the expected frequency content of the resulting signal, would it work best on big or small bubbles, would it also work on lipid bubbles, etc.

Line 30: “In the later years […] I might suggest adding a reference here (like a review article or something, perhaps a Klibanov one?)

Line 33: “less resonant”. I think you’ll need to clarify this. What is a less resonant bubble? You mean it’s resonant oscillation magnitude is smaller than that of a resonant lipid-coated bubble of the same size and shell properties? I’m not sure that’s such an easy comparison but I do see your point. Please clarify

Line 35: “Compared to ultrasound imaging of thin-shelled MBs, higher mechanical indices are typically required to drive thick-shelled MBs into nonlinear oscillations.”.

• Please provide a reference

• It’s generally understood that thick-shelled bubbles need to break to elicit sufficient nonlinear oscillations (i.e. destructive nonlinear behaviour). Is this what the authors are referring to?

Line 38: “[…] while supressing harmonic components from soft tissues.” These pulses are also designed to supress fundamental signal from soft tissues, too. Please modify sentence to account for this.

Microbubble preparation: Were these measured by Coulter Counter also? Microscopy methods will not capture bubbles below 1 um (biased result). How confident are the authors in this size distribution (Fig 1)? A concentration of 1e8 is quite low, which might be a symptom of missing the smaller bubbles.

In-vitro testing: Clearly the comparison made between the novel pulse sequence and traditional pulse-inversion isn’t valid, since the MI’s are not the same and neither are the frequencies! I’m a little confused: If the authors used PI at an MI equal to that of their new sequence (MI=1), they surely would obtain usable contrast-images. Is this true? For example, the ‘new method, 2 pulses’ is actually sort of similar to AM (since there is no intervening pulses), but taken at MI=1.

In fact, this novel scheme works essentially as amplitude modulation. The idea is that the intervening pulses either move or distort the bubbles (but not tissue), and then upon subtraction of two pulses, you are left with ‘nonlinear fundamental’ signal – that is signal at the fundamental (which is linear) due to the presence or absence of a target. I’d like to see how this new sequence compares with traditional AM (which is almost the same as ‘new method, 2 pulses’).

Reviewer #2: In this manuscript entitled “A multi-pulse ultrasound technique for imaging of thick

shelled microbubbles demonstrated in vitro and in vivo”, the authors report a multi-pulse subtraction imaging technique, with interleaved manipulation pulses, to image thick shelled polymer microbubbles in vitro in a flow loop and in vivo in a rabbit model of inflammation. Their in vitro experiments are conducted using very low concentration of MB in a commercially available flow phantom. Their in vitro results indicate that their non-destructive subtraction imaging scheme, which, on responding microbubbles, can result in a CTR of 23dB, which is good. This is attributed to a translation of the MB of up ~4um, which is roughly the diameter of one MB. The manipulating pulses can theoretically be modified (pressure, length, frequency) or even completely removed, which they did for their in vivo experiments.

In vivo, they used a modified version of their imaging approach (without manipulation pulse) for the molecular imaging of ICAM1-targeted MB using and anti-ICAM1 decorated MB using biotin streptavidin conjugation, which is an adequate conjugation scheme for a proof of principle study. They demonstrate that MB can be detected at 24h when MB are injected at 2h post injury and 72h when injected at 24h post injury but not with control MB. This is supported by histological fluorescence imaging in N=1 rabbit per condition.

Overall, this reports contains some interesting data but (1) suffers from unsupported statements and inadequate experiments in vitro and (2) lacks adequate quantitative analysis in vivo (essentially N=1 and non-analysis of longitudinal data)

MAJOR FLAWS

1- For instance, since their imaging application is targeted MB, it is not acceptable to only characterize MB in flowing conditions in vitro. Especially since their approach relies on MB translation caused by radiation force. Indeed, one has to wonder if that translation will be hindered once MBs are attached. In vitro experiments with targeted MB should be performed. How many MB will respond ? Will they roll with successive pulses? Or detach ?

2- Also, in the first part of the manuscript, the authors optimize the manipulation pulses in vitro but end up using a different imaging scheme in vivo without manipulation pulses. It would be relevant to repeat the in vitro experiments with the conditions used in vivo (no manipulation) to allow comparison and understand the importance of the manipulation pulses.

3- Subtraction methods are known to be sensitive to motion. Yet their images (cf Fig 7) clearly do not reflect this reality, which is clearly visible in the in vivo supplementary videos. This limitation should be disclosed and discussed. Why is there so much movement artifact with rabbit 3 ? How does it affect the CTR ?

4- Why are only a fraction of MB responding to this imaging scheme (Figure 5)? Please quantify that. The reviewer wonders if it is fair to quantify only responding MB in table 1. Since you report results on 6-10 MB in table 1, how did you choose which MBs to analyze ? There are clearly more than 10 MB in Fig4. How did you account for the MBs that do not give signal with your method (cf Fig 5 A and B)?

5- I have never seen PI harmonic imaging with such poor tissue cancellation performance (cf Figure 4). Was this contrast package the commercial package from GE or was it home-made ? Please specify in the methods. Also, an additional in vitro image with a lipid MB, perhaps in supplementary data, would be appropriate to convince the readers of the proper coding of the PI sequence.

6- Moreover, I strongly encourage the authors to use and report the same dynamic range in all images. This is an important parameter that largely affects the visual rendering of CTR

7- Although this is a proof of principle study, reporting data with N=1 is never a good approach. Perhaps one way to bring more robustness to the study would be to analyze the images longitudinally at different time points (48h vs 72h), which you state you have performed in the methods, but are not reporting or quantifying. For Rabbit 1, you could even do 24h, 48h and 72h. L461: replace “days” by “48 hours in one animal”.

8- Is it possible that your inflammation model has some variability (hard to know with N=1) ? Could Rabbit 3 have less inflammation than rabbit 2 explaining the change in brightness and histology ? Could you analyse ICAM1 expression using histology ?

9- The selections of the analyzed sections in the histology analyses need to be shown. It is unclear how the 11 areas in the inflamed tissues and 7 areas in control muscle were selected.

10- How were in vivo CTR values computed ? How did you choose MB signal vs tissue signal ? In which frame in the video ? Please describe the method and report mean, standard deviation, #MB analyzed and how they were chosen.

11- It would be interesting to stain for endothelium to confirm the compartment of the microbubbles after 48h in vivo. Are they in the vasculature ? Uptaken by leucocytes ?

MINOR

You mention imaging during MB injection (L204) but don’t show images. Since two B-mode images are used to compute the contrast image, using one for B_mode would allow perfect coregistration and allow anatomical correspondence in Figure 7.

Please verify the excitation wavelength for your fluorophores L217-L218

Figure 7 : Perhaps using the first imaging pulse as the B-mode would help coregister B-mode and contrast images. Right now you are selecting an arbitrary frame from the cineloop than is not in the same position as the B-mode

Figure 7E : Correct label on the right side : It should be 24h correct ?

Figure 8 : Use arrows to indicate the different tissues (including the puss filled region and normal muscle tissue) and indicate features described in the text (inflammatory cells, dead leukocytes, cellular debris, granulocytes). Explain why these two sections were selected.

Reviewer #3: This manuscript describes a contrast pulse strategy to interlace 2 modifications pulses between 2 identical imaging pulses for the specific application of thick-shell targeted microbubbles. While such method can’t pick-up non-linear echoes like Pulse Inversion, it should be sensitive to agent motion and disruption. The ability to detect bounded microbubbles in-vivo 22 hours after injection is impressive. My main concern is that the authors use identical pulses for the imaging and modification components in this manuscript which is very close to a Doppler sequence (identical to a Doppler sequence if the PRF is constant). The in-vivo section uses 2 pulses which is identical to a (2-pulse) Doppler sequence. Doppler does work at detecting bubbles at high MI regimes but is not novel. While I understand the research limitation of the clinical system, my suggestion to the authors is to either design a simple single transducer in-vitro experiment demonstrating the advantage of the intermediate pulses (e.g. much longer pulse, different frequency, etc.) to distinguish the method from Doppler, or rephrase the manuscript as using Doppler for targeted agent which by itself could be novel. In the later case, measuring the Doppler spectrum of the targeted agents and comparing it to tissue would be relevant.

The introduction is also missing significant background on existing pulse sequences. Amplitude modulation in particular is relevant to this work as its application has been demonstrated to gas vesicle which closely resemble thick shell microbubbles. It is also routinely available on clinical ultrasound systems.

Comments:

(introduction) ‘Thick-shelled MBs can have improved circulation half-lives, incorporate larger amounts of drugs for enhanced drug delivery or facilitate targeting for use in molecular ultrasound imaging. However, methods for robust imaging of thick-shelled MBs are currently not available.” Comment: What does the author consider thick shell? Would Sonazoid count as a stiff shell (compared to Sonovue/Definity)?

(introduction) The introduction does not provide information about the proposed novel imaging technique. Also, the background on general pulse sequence literature is missing from the introduction.

(p.5) ‘MBs used in the current experiments were not well imaged by conventional harmonic imaging schemes based on pulse inversion’ Comment: What about amplitude modulation (AM)? In particular, one would suspect that AM would perform well for thick shell agents due to their buckling dynamics (see Marmottant model)

(p.6) The description of the pulse sequence comes really late in the manuscript and is hidden with other experimental details. I would move it closer to the introduction.

(p.6) ‘first and the last pulse were identical, and were used for imaging, whereas one or several intermediate pulses could be included for additional manipulation of the MB’. Comment: If the first and the last pulses are identical, then the pulse sequence is not extracting any nonlinear echo information but instead will pick-up either motion, bubble disruption or partial bubble disruption. Furthermore, here, both the imaging and intermediate pulses are identical which means the authors are simply using a Doppler sequence (in particular if the PRF is constant). This would need clarification in the manuscript.

(p.7 l.137) ‘and acquisition schemes with MI of 0.7 and 1.0’. Comment: While I expect hard shell bubbles to be more resilient to pressure, those are very high MI when compared to conventional contrast imaging. For traditional non-destructive imaging the norm is MI~0.06-0.1. The introduction made is sound like the contrast sequence needs to be non-destructive. This should be clarified in the manuscript.

(p.7) ‘The ultrasound transducer was fixed with a clamp and MBs were imaged both during low speed flow and when the pump had stopped and the MBs were stationary.’ Comment: Were results computed for each condition or combined? Flow introduce motion between pulses and yield a significant increase in the CTR which may shadow other effects from the intermediate pulses.

(p.10) ‘A frequency of 9.5 MHz and MI of 1 was used in the contrast images, and 12 MHz and MI of 1 was used to acquire the B-mode images. In order to minimize tissue signal and increase the framerate, the in vivo recordings were done without any intermediate pulses, hence only the two imaging pulses were transmitted along each beam line. Comment: In this case, the contrast sequence is simply a 2 pulse Doppler sequence. Doppler sequence do work at imaging bubbles, but only if they are moving are destroyed by the imaging pulse. Doppler was actually the original detection scheme for microbubbles (before PI and AM were invented), and has been documented in literature. I recommend that the authors discuss this in the introduction.

(p.12) ‘When comparing the amount of MBs in the image in Figs 5A and 5B, it is apparent that not all MBs become visible with this method’. Comment: Could this indicate that the high MI has detrimental effect on the thick shell microbubbles or that microbubble coalesce into larger bubbles that become easier to detect?

(p.13) ‘The backscattered signals from 6 to 10 MBs in each image were analyzed, and the detected mean delay between the first and last imaging pulse along the scanline through the MB and the mean CTR are presented in Table 1.’ Comment: This CTR algorithm seems biased. Bubbles are supposed to be everywhere in the flow channel; a more robust metric would be to compare larger ROI in the tube and tissue.

(p.14) ‘The CTR of circulating MBs immediate after injection was 15-20 dB.’ Comment: It would be relevant to create a new figure with the Contrast image a several time stamps after the injection in order to justify the 15-20dB claim.

(Discussion) It should be discussed why the authors didn’t consider Amplitude modulation, which is the most commonly used contrast pulse sequence on clinical systems. It is also worth noting that gas vesicles (nanobubbles with thick shells) were successfully imaged with AM due to their buckling dynamics [reference: Acoustic Behavior of Halobacterium salinarum Gas Vesicles in the High-Frequency Range: Experiments and Modeling, E. Cherin et al., UMB, Vol 43, Issue 5, pp.1016-1030] and is therefore a promising candidate for their thick shell microbubbles.

(p.19 l.428) ‘Tissue motion may especially be a problem in the abdominal and thoracic areas, where breathing and the beating heart may cause strong artefacts if no motion compensation is included in the implementation. In the current work, a relatively stationary organ was imaged (the thigh).’ Comment: This is precisely the same limitation of Doppler imaging. Doppler mitigates this issue by using a wall-filter which rejects low Doppler frequencies. I recommend to measure the Doppler spectrum from these microbubbles and compare it to the in-vivo Doppler spectrum of tissue. One would expect the microbubble spectrum bandwidth to broaden compared to tissue Doppler due to signal decorrelation.

(p.19 l.437) ‘We have not been able to verify the presence of such effects in the current work’. Comment: That being said, it is unlikely that the acoustic pulses moved the bounded bubbles in the in-vivo experiment.

Editorial

(p.5 l.87) ‘[…] has been described earlier’ Comment: Please rephrase as ‘has been described by Cavalieri et al […]’

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Reviewer #1: No

Reviewer #2: No

Reviewer #3: No

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PLoS One. 2022 Nov 3;17(11):e0276292. doi: 10.1371/journal.pone.0276292.r002

Author response to Decision Letter 0

8 Apr 2022

We thank for the thorough review done by all three reviewers. We have addressed all the comments and answered the remarks in the attached document. Changes made in the manuscript are commented with reference to line numbers in the revised version of the manuscript.

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Submitted filename: Response to reviewers.docx

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PLoS One. doi: 10.1371/journal.pone.0276292.r003

Decision Letter 1

Guy Cloutier

2 May 2022

PONE-D-21-28949R1A multi-pulse ultrasound technique for imaging of thick-shelled microbubbles demonstrated in vitro and in vivoPLOS ONE

Dear Dr. Berg,

Please submit your revised manuscript by Jun 16 2022 11:59PM. If you will need more time than this to complete your revisions, please reply to this message or contact the journal office at plosone@plos.org. When you're ready to submit your revision, log on to https://www.editorialmanager.com/pone/ and select the 'Submissions Needing Revision' folder to locate your manuscript file.

Please include the following items when submitting your revised manuscript:

A rebuttal letter that responds to each point raised by the academic editor and reviewer(s). You should upload this letter as a separate file labeled 'Response to Reviewers'.
A marked-up copy of your manuscript that highlights changes made to the original version. You should upload this as a separate file labeled 'Revised Manuscript with Track Changes'.
An unmarked version of your revised paper without tracked changes. You should upload this as a separate file labeled 'Manuscript'.

We look forward to receiving your revised manuscript.

Kind regards,

Guy Cloutier, Ph.D.

Academic Editor

PLOS ONE

Additional Editor Comments:

Unfortunately, one reviewer still believes that another round of revision is necessary. He is insisting on the use of amplitude modulation in new sets of experiments for comparison. I will wait for your new responses and changes before finalizing the decision.

Sincerely,

Dr Cloutier

Academic Editor

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Reviewers' comments:

Reviewer's Responses to Questions

Comments to the Author

1. If the authors have adequately addressed your comments raised in a previous round of review and you feel that this manuscript is now acceptable for publication, you may indicate that here to bypass the “Comments to the Author” section, enter your conflict of interest statement in the “Confidential to Editor” section, and submit your "Accept" recommendation.

Reviewer #1: All comments have been addressed

Reviewer #2: All comments have been addressed

Reviewer #3: (No Response)

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2. Is the manuscript technically sound, and do the data support the conclusions?

Reviewer #1: (No Response)

Reviewer #2: Yes

Reviewer #3: Partly

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3. Has the statistical analysis been performed appropriately and rigorously?

Reviewer #1: N/A

Reviewer #2: N/A

Reviewer #3: Yes

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4. Have the authors made all data underlying the findings in their manuscript fully available?

Reviewer #1: (No Response)

Reviewer #2: Yes

Reviewer #3: Yes

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5. Is the manuscript presented in an intelligible fashion and written in standard English?

Reviewer #1: Yes

Reviewer #2: Yes

Reviewer #3: Yes

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6. Review Comments to the Author

Reviewer #1: The authors do an adequate job at replying to the reviewer's commentary. I will just clarify one thing: The authors state that AM relies on nonlinear behaviour and their proposed method relies on radiation force and is thus mechanistically different. Typically, when we refer to nonlinear bubble vibrations we are referring to their nonlinear content (e.g. subH), and I agree that in this context (thin versus thick shell), the stated difference above is reasonable since thick-shelled bubbles don't vibrate very easily. However, the power of AM comes from the 'nonlinear' signal remaining in the fundamental band due, not specifically to nonlinear vibrations, but to a non-proportional response to two input pressures (at the fundamental - i.e. linear band). This is due to the fact that a given bubble may not vibrate at all at one pressure amplitude A, and then vibrate with a very large amplitude at a transmit pressure 2A - more than just a proportional amount. This is not a nonlinear oscillation issue, it's a lack/presence of microbubble issue. In this sense, the proposed technique works very, very similarly (if not identically) to AM, as was mentioned by both reviewers.

Reviewer #2: Thank you for addressing most of my comments.

- I am not certain I understand your response to Question 1 about additional in vitro experiments:

"Unfortunately, we do not have the needed infrastructure for conducting such experiments at this point."

What infrastructure are you referring to ? Targeting MB in vitro is not that difficult. For example you can look at PMID: 19411212.

- I also think that you could do a better job at describing and quantifying your CTR in vivo results L351-356. Since you have 172 + 205 MB you have analyzed at different time points, it would be interesting to describe these results in more detail using a figure.

There are still a few typos that may require your attention

L252 : Grammar : satisfactory is an adjective. In this phrasing your should use an adverb such as "adequately" ? or perhaps write : "was low, as expected"

L313: Grammar : ... in Table 1. "Additionally to the results in Table 1," RF data from images...

L319 : In table1 : I believe that a period is missing for MI=0.7 3HP standard deviation

L343: Please add : 16.8dB "for flowing MB"

L353 : Please add : 13.3 dB "for static MB"

L378 : you say 30 dB but the color bar is suggesting 35 dB. Please verify.

L383 : Is is common to display DAPI in blue. I understand this is pseudo colouring but it is confusing when looking at the image. If possible please use blue for DAPI.

L516 : Grammar : please add a period and cut in two sentences : ... [14,15]. This is also supported by...

L518 : replace "towards" with "of"

Reviewer #3: I would like to thank the author for revising the manuscript. However, I still believe it would be worthwhile to properly investigate the advantage of the manipulation pulse both in-vitro and in-vivo. While the modulation has theoretical advantages over power Doppler, this is not quite demonstrated in the manuscript. I also think it would be worthwhile to compare the proposed sequence to Amplitude modulation in a flow phantom setup (as done for PI).

(l. 326) “However, due to motion, and flashing artefacts when using the intermediate manipulation pulses in vivo, the in vivo results were recorded without manipulation pulses at MI=1.”

Comment: The addition of the manipulation pulse will decrease the pulse repetition frequency between the imaging pulses, but should also increase the bubble displacement and improve detection from background tissue. Similar to Power Doppler, the detectability of the microbubbles will be tied to the amount of displacement detected relative to surrounding tissue displacement.

(From reviewer response) “Gas vesicles typically have very thin protein shells (approx. 2nm or less) and the vesicle itself has a cylindrical dimension with diam of 50-100 nm and length of 100-1000 nm. The PVA bubble has a thick polymer shell of approx. 200 nm. Our GE Vivid scanner did not have an implementation of AM and there is no reason to believe that AM will work adequately for thick-shelled PVA bubbles. The buckling-effect and compression-only behavior for example seen with SonoVue (having a lipid shell of approx. 2 nm) is unlikely with a thick polymer shell of approx. 200 nm.”

Comment: While it is true that gas vesicle differs from thick shell microbubbles, there is experimental and theoretical evidence in literature that thick shell microbubbles will buckle through reversible collapse. In particular see [P. Marmottant et al. Buckling resistance of solid shell bubbles under ultrasound, JASA, 129 (3), 2011]. As the CTR achieved by AM is strongly affected by the buckling dynamics of bubbles, there is reason to believe that AM will perform much better than PI [ref: C. Tremblay-Darveau, IEEE Trans UFFC, 65 (8), 2018]. It seem highly relevant, and the authors should ideally perform an in-vitro validation of their technique to AM (which should be readily available on most clinical systems) even if done on a different clinical system.

(From reviewer response) “ One might argue that the technique has similarities with conventional Doppler but Doppler techniques are made for imaging an existing flow, not to induce a translation due to ultrasound radiation force as with our method. That is why we want to use imaging pulses tailored to the organ size and depth of interest and intermediate manipulation pulses that are tailored to maximize the radiation force on the PVA bubbles. Also, Doppler makes use of transmit pulses consisting of several pulse oscillations (resulting in reduced range resolution) and several pulse packets needs to be transmitted in each beam direction for spectral analysis (to construct an adequate clutter filter for suppressing strong, slow moving scatterers and to estimate blood velocities).”

Comment: I agree that pulses used for CW and PW are typically quite long, but Colour or Power Doppler pulses are typically of length comparable to that of Contrast pulses. If the manipulation pulse is taken out (as done in the in-vivo section), then the pulse sequence is equivalent to a N=2 ensemble power Doppler sequence (and is not novel). It would be worthwhile to try the power Doppler mode on the GE machine (using the lowest PRF) and see how it compares to the proposed pulse sequence. It is also not clear to me why the manipulation pulse is taken out for the in-vivo imaging. The authors claim that using the manipulation pulse causes too much tissue motion clutter (l.326), but clutter is a fundamental limitation of any Doppler system. Does this mean that the implementation of the manipulation pulse is not practical in-vivo? I understand the theoretical advantages of using a long burst for the manipulation pulse, but this is not demonstrated in-vivo in the manuscript.

(From reviewer response) “The aim of the current paper was not to describe various contrast pulse sequences and reference 3 provides a good overview of such sequences”

Comment: The reader should be able to infer from the introduction where the authors work sits within the existing literature. For instance, is there any literature work describing how PI/AM/AMPI do not work well with hard shell microbubble? Any theorical work on solid shell microbubble that could be use to justify the pulse sequence (Marmottant’s paper on stiff shell buckling dynamic seems relevant)? If the goal of the modulation pulse is to displace microbubble, some citation of papers describing acoustic radiation forces on bubbles is also relevant.

(From reviewer response) “ (p.19 l.437) ‘We have not been able to verify the presence of such effects in the current work’. -> That being said, it is unlikely that the acoustic pulses moved the bounded bubbles in the in-vivo experiment. -> We agree. Knowing exactly what the cause of the detection signal from each MB in an in-vivo situation is impossible (except for maybe in intra-vital microscopy at very high magnification and frame rate). In vitro, it is possible to study single bubble behavior, as has been done in Kohtapalli at al (2015).”

Comment: This limitation should be described in the Discussion section, including possible hypothesis from existing literature.

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Reviewer #1: No

Reviewer #2: No

Reviewer #3: No

PLoS One. 2022 Nov 3;17(11):e0276292. doi: 10.1371/journal.pone.0276292.r004

Author response to Decision Letter 1

29 Jun 2022

We thank for the second round of thorough review done by all three reviewers. We have addressed the comments and answered the remarks in the document included in the submission. Changes made in the manuscript are commented with reference to line numbers in the revised version of the manuscript. Images showing a comparison of the novel method to both Amplitude Modulation and Pulse Inversion is now included in the manuscript, as requested.

Attachment

Submitted filename: Respons to Reviewers_2.review.pdf

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PLoS One. doi: 10.1371/journal.pone.0276292.r005

Decision Letter 2

Joseph Donlan

4 Aug 2022

PONE-D-21-28949R2A multi-pulse ultrasound technique for imaging of thick-shelled microbubbles demonstrated in vitro and in vivoPLOS ONE

Dear Dr. Berg,

Your manuscript has been assessed again by the three reviewers from the previous rounds, who are satisfied that their concerns have been addressed. However, when carrying out our editorial assessment of this manuscript, we noted that your data availability statement states that some underlying data cannot be shared because of existing non-disclosure agreements. PLOS journals require authors to make all data necessary to replicate their study’s findings publicly available without restriction at the time of publication - restrictions on data availability because of commercial reasons such as non-disclosure agreements do not comply with this policy (see full details of the PLOS data policy at https://journals.plos.org/plosone/s/data-availability). To comply with the PLOS data policy, please ensure data underlying all of your findings are made available, outlining in your data availability statement how the data can be accessed. Alternatively, if it is not possible to make all data available, please remove the parts of the manuscript for which the raw supporting data cannot be provided and include a detailed justification in your response-to-reviewers that outlines which data were removed and if/how your conclusions are affected. We will not be able to proceed further with your manuscript until these concerns are addressed.

Please submit your revised manuscript by Sep 17 2022 11:59PM. If you will need more time than this to complete your revisions, please reply to this message or contact the journal office at plosone@plos.org. When you're ready to submit your revision, log on to https://www.editorialmanager.com/pone/ and select the 'Submissions Needing Revision' folder to locate your manuscript file.

Please include the following items when submitting your revised manuscript:

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An unmarked version of your revised paper without tracked changes. You should upload this as a separate file labeled 'Manuscript'.

We look forward to receiving your revised manuscript.

Kind regards,

Dr Joseph Donlan

Senior Editor

PLOS ONE

Journal Requirements:

Please review your reference list to ensure that it is complete and correct. If you have cited papers that have been retracted, please include the rationale for doing so in the manuscript text, or remove these references and replace them with relevant current references. Any changes to the reference list should be mentioned in the rebuttal letter that accompanies your revised manuscript. If you need to cite a retracted article, indicate the article’s retracted status in the References list and also include a citation and full reference for the retraction notice.

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Reviewers' comments:

Reviewer's Responses to Questions

Comments to the Author

Reviewer #1: All comments have been addressed

Reviewer #2: All comments have been addressed

Reviewer #3: All comments have been addressed

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2. Is the manuscript technically sound, and do the data support the conclusions?

Reviewer #1: (No Response)

Reviewer #2: Yes

Reviewer #3: Yes

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3. Has the statistical analysis been performed appropriately and rigorously?

Reviewer #1: (No Response)

Reviewer #2: Yes

Reviewer #3: Yes

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4. Have the authors made all data underlying the findings in their manuscript fully available?

Reviewer #1: (No Response)

Reviewer #2: Yes

Reviewer #3: Yes

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5. Is the manuscript presented in an intelligible fashion and written in standard English?

Reviewer #1: (No Response)

Reviewer #2: Yes

Reviewer #3: Yes

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6. Review Comments to the Author

Reviewer #1: (No Response)

Reviewer #2: Thank you for addressing my concerns.

In the end, the quality of your in vivo results convinced me that this should be published.

Reviewer #3: Thank you for including the new AM data to the manuscript and for the detailed responses. I have no additional comment.

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7. PLOS authors have the option to publish the peer review history of their article (what does this mean?). If published, this will include your full peer review and any attached files.

If you choose “no”, your identity will remain anonymous but your review may still be made public.

Do you want your identity to be public for this peer review? For information about this choice, including consent withdrawal, please see our Privacy Policy.

Reviewer #1: No

Reviewer #2: No

Reviewer #3: No

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PLoS One. 2022 Nov 3;17(11):e0276292. doi: 10.1371/journal.pone.0276292.r006

Author response to Decision Letter 2

7 Sep 2022

After communication with representatives from GE Healthcare, we were allowed to also share the IQ-data which in the first versions were not included in the original submission. This is the data that forms the basis of Table 1 and Figure 5 in the manuscript. The IQ data is included in the data repository as .mat-files.

The manuscript is re-submitted unchanged.

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PLoS One. doi: 10.1371/journal.pone.0276292.r007

Decision Letter 3

Joseph Donlan

5 Oct 2022

A multi-pulse ultrasound technique for imaging of thick-shelled microbubbles demonstrated in vitro and in vivo

PONE-D-21-28949R3

Dear Dr. Berg,

We’re pleased to inform you that your manuscript has been judged scientifically suitable for publication and will be formally accepted for publication once it meets all outstanding technical requirements.

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Kind regards,

Dr Joseph Donlan

Senior Editor

PLOS ONE

Additional Editor Comments (optional):

Reviewers' comments:

PLoS One. doi: 10.1371/journal.pone.0276292.r008

Acceptance letter

Joseph Donlan

12 Oct 2022

PONE-D-21-28949R3

A multi-pulse ultrasound technique for imaging of thick-shelled microbubbles demonstrated in vitro and in vivo

Dear Dr. Berg:

I'm pleased to inform you that your manuscript has been deemed suitable for publication in PLOS ONE. Congratulations! Your manuscript is now with our production department.

If your institution or institutions have a press office, please let them know about your upcoming paper now to help maximize its impact. If they'll be preparing press materials, please inform our press team within the next 48 hours. Your manuscript will remain under strict press embargo until 2 pm Eastern Time on the date of publication. For more information please contact onepress@plos.org.

If we can help with anything else, please email us at plosone@plos.org.

Thank you for submitting your work to PLOS ONE and supporting open access.

Kind regards,

PLOS ONE Editorial Office Staff

on behalf of

Dr Joseph Donlan

Staff Editor

PLOS ONE

Associated Data

This section collects any data citations, data availability statements, or supplementary materials included in this article.

Supplementary Materials

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Data Availability Statement

Underlying data is fully available at the following repository: https://doi.org/10.11582/2021.00072.

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PERMALINK

A multi-pulse ultrasound technique for imaging of thick-shelled microbubbles demonstrated in vitro and in vivo

Sigrid Berg

Siv Eggen

Kenneth Caidahl

Lars Dähne

Rune Hansen

Roles

Abstract

Introduction

Materials and methods

Microbubble preparation

Fig 1. Microscopy image of a typical sample of the polymeric MBs (A), and size distribution calculated based on automatic counting of optical microscopy images (B).

In-vitro ultrasound imaging

Fig 2. Illustration of the proposed multi-pulse imaging technique.

In-vivo experiments

Chemicals

Animal preparation and experimental procedures

Fig 3. Illustration of timeline and experimental procedures in the in-vivo experiments.

Molecular ultrasound imaging

Histology and microscopy imaging

Results

In vitro ultrasound imaging

Fig 4. Ultrasound images of MBs in a tissue mimicking flow phantom.

Fig 5. A representative example of the signal from a MB in a tube in a tissue mimicking phantom.

Table 1. Mean delay and CTR ratio in images recorded at 8 and 9.5 MHz.

In-vivo ultrasound imaging

Fig 6. Ultrasound B-mode images from Rabbit 1.

Fig 7. Examples of B-mode (A, C and E) and contrast enhanced (B, D and F) images from the three rabbits included in the study.

Muscle tissue necropsy

Ex vivo analysis

Fig 8. Example of HES (A) and corresponding DAPI mounted (B-D) tissue sections from the inflamed tissue from the rabbit that was injected with strep-tritc-MBs with anti-ICAM-1 2h after inflammation was induced and sacrificed 70h later.

Fig 9. Number of MBs found in a 1mm x 1 mm region of interest (ROI) in 25 μm thick tissue sections from inflamed (dark colors) and healthy (light colors) muscle tissue from rabbits.

Animal health

Discussion

Ultrasound imaging technique

The inflammatory reaction

Molecular ultrasound imaging

Conclusion

Supporting information

Acknowledgments

Data Availability

Funding Statement

References

Decision Letter 0

Guy Cloutier

Roles

Author response to Decision Letter 0

Decision Letter 1

Guy Cloutier

Roles

Author response to Decision Letter 1

Decision Letter 2

Joseph Donlan

Roles

Author response to Decision Letter 2

Decision Letter 3

Joseph Donlan

Roles

Acceptance letter

Joseph Donlan

Roles

Associated Data

Supplementary Materials

Data Availability Statement

ACTIONS

PERMALINK

RESOURCES

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